Conductive sensors and their use in diagnostic assays

ABSTRACT

A conductive sensor and its use in a diagnostic assay are disclosed. The miniaturized conductive sensor, utilizing a conducting polymer, is used in a diagnostic device to determine the presence or concentration of a predetermined analyte in a liquid test sample, wherein the predetermined analyte, like glucose, is assayed by an oxidase interaction. The interaction between the oxidase and a small amount of the predetermined analyte in the test sample generates, either directly or indirectly, a dopant compound in a reaction zone of the conductive sensor. The dopant compound then migrates to the detection zone of the conductive sensor of the diagnostic device to oxidize the conducting polymer and convert the conducting polymer from an insulating form to a conducting form. The resulting increase in conductivity of the conducting polymer is measured, then the conductivity increase is correlated to the concentration of the predetermined analyte in the test sample.

This is a continuation, of application Ser. No. 554,393, filed Jul. 19,1990, now abandoned.

FIELD OF INVENTION

The present invention relates to a method of determining the presence orconcentration of a predetermined analyte in a test sample with adiagnostic device including a miniaturized conductive sensor. Theconductive sensor comprising a reaction zone including an enzyme thatselectively interacts with the predetermined analyte and a detectionzone including a layer or film of a conducting polymer and amicroelectrode assembly. More particularly, a diagnostic device isemployed to selectively assay a test sample for the presence orconcentration of a specific predetermined analyte by measuring thechange in conductivity of a layer of conducting polymer present in thedetection zone of the conductive sensor. The conductivity change of thelayer of conducting polymer results from the generation of a dopantcompound in the reaction zone of the sensor due, either directly orindirectly, to an enzymatic interaction with the predetermined analyte.The dopant compound then migrates from the reaction zone of the sensorto the detection zone of the conductive sensor to dope the layer ofconducting polymer and change the conductivity of the polymer. Forexample, a dopant compound, molecular iodine, is formed in the reactionzone in a reaction between iodide ions, a peroxidase enzyme or amolybdenum(VI) transition metal catalyst; and the hydrogen peroxideformed from a glucose oxidase interaction with glucose. The dopantcompound then migrates, or diffuses, from the reaction zone to dope,oxidatively, the layer of conducting polymer in the detection zone.Surprisingly and unexpectedly, the configuration of the conductivesensor is such that only a small fraction of the glucose in the testsample, such as about 1% or less, is enzymatically converted to generatethe dopant compound. Accordingly, oxygen limitations in the enzymeinteraction are avoided. However, the amount of a generated dopantcompound is sufficient to change the conductivity of the conductingpolymer layer in the conductive sensor and allow an accurateconductivity measurement that can be correlated to the amount of theglucose in the test sample.

BACKGROUND OF THE INVENTION AND PRIOR ART

Reagent-impregnated test strips are commercially-available for a varietyof predetermined analytes, such as glucose. These test strips areaccurate, economical and relatively easy-to-use by an individual athome. The availability of such test strips has played a major role inthe decentralization of the diagnostic market and has played anespecially critical role in the development of home blood glucosemonitoring that enables better glycemic control for diabeticindividuals. In addition, other analyte detection devices designed forhome use and that utilize amperometric, electrochemical detection of thepredetermined analyte also are available. However, two importantlimitations remain in currently available technologies for homemonitoring of a predetermined analyte, including the sample volume andthe length of time that the assay requires. Generally, a test samplevolume of between 5 μL and 20 μL (microliters) is required to perform anassay. The test sample is obtained by a finger puncture that producesfresh capillary blood. This is a relatively painful procedure, and thediscomfort involved affects the willingness of an individual to performassays as often as medically useful. In addition, currently-availableassays require between 30 seconds and two minutes to complete.

Investigators have therefore shown an intense interest in developingelectrochemical sensors that can be miniaturized, and therefore eitherbe implanted subcutaneously or require a much smaller sample volume.Numerous approaches have been tried, including amperometric sensorsbased upon fiber electrodes and potentiometric sensors manufacturedusing techniques established in the semi-conductor industry. Incontrast, the present invention measures a conductivity change in a thinpolymer film, and accordingly has solved many of the problems inherentin miniaturized electrochemical sensors. Most importantly, anelectrochemical sensor of the present invention can be manufactured in aplanar format, using semi-conductor technologies, to provide aneconomical, disposable sensing element. Also, the sensor can be madesufficiently small such that a sample volume of 1 μL or less can beassayed. The sensitivity of the detection method also has solved thegeneral oxygen limitation problem frequently observed in electrochemicalsensors that utilize oxidase enzymes. Therefore, in accordance with animportant feature of the present invention, an electrochemical sensorthat can be miniaturized; that provides rapid and accurate assays; thatovercomes oxygen limitation problems; that is free of interferencesattributed to common constituents in the test sample; and that can beproduced economically is achieved.

In general, a diagnostic device of the present invention comprises areaction zone wherein the predetermined analyte of interest interactswith a suitable oxidase enzyme to generate, either directly orindirectly, a dopant compound. The dopant compound is capable ofoxidizing a conducting polymer to change the conductivity of thepolymer. In addition, the diagnostic device further comprises adetection zone including a film or layer of conducting polymer and amicroelectrode assembly. The change in conductivity of the conductingpolymer layer as a result of the dopant compound is detected or measuredby the microelectrode assembly. The change in conductivity then iscorrelated to the amount of predetermined analyte in the test sample. Aswill be demonstrated more fully hereinafter, an economical andreproducible conductive sensor, useful in assaying for a predeterminedanalyte that responds to oxidase chemistry, has been provided. Theconductive sensor utilizes the properties of conducting polymers, avoidsoxygen limitation problems in the normal concentration ranges of thepredetermined analyte; can be manufactured by well-known semiconductorprocessing techniques; and does not rely upon a chemical reactionoccurring at the microelectrode assembly.

Accordingly, one important feature of the conductive sensor of thepresent invention is the conducting polymer included in the detectionzone of the conductive sensor. Several organic conducting polymers, suchas polyacetylene, polypyrrole and polythiophene, are known. The organicconducting polymers have several potential applications in the fields ofbatteries, display devices, corrosion prevention in metals andsemiconductors and in microelectronic devices such as diodes,transistors, sensors, light emitting devices and energy conversion andstorage elements. However, organic conducting polymers possess severallimitations that have hindered the use of organic conducting polymers inconductive sensors. In general, a conducting polymer useful in theconductive sensor of the present invention should display a sufficientlyhigh conductivity for detectable and accurate measurements, and shouldbe capable of being processed, reproducibly, into thin, uniform films.Although several conducting polymers possess one of these two necessaryproperties, only a limited number of conducting polymers possess bothnecessary properties.

For example, polyacetylene acts as an insulator exhibitingconductivities in the range of 10⁻¹⁰ S/cm to 10⁻¹³ S/cm (Siemens percentimeter). A conductivity in this range corresponds to theconductivity of known insulators, such as glass and DNA. However,polyacetylene can be doped using a variety of oxidizing or reducingagents, such as antimony pentafluoride, the halogens, arsenicpentafluoride or aluminum chloride. By doping, polyacetylene isconverted into a highly-conducting polymer exhibiting a conductivity ofapproximately 10³ S/cm, corresponding to the conductivity of metals suchas bismuth.

However, polyacetylene suffers from the drawbacks of extreme instabilityin air and a sharp drop in conductivity when an alkyl or othersubstituent group is introduced into the polymer. Accordingly, theinstability of polyacetylene in the presence of oxygen, and theinability of substituted polyacetylenes to maintain a high conductivity,generally makes polyacetylene unsuitable as the conducting polymer in aconductive analyte sensor. As will be discussed more fully hereinafter,conducting polymers including alkyl or other substituent groups usuallypossess physical and mechanical properties making the substitutedconducting polymer more easily processible into conducting polymericfilms than the corresponding unsubstituted conducting polymer.Therefore, in the manufacture of conductive sensors it is desirable touse an easy-to-process conducting polymer, such as a conducting polymerthat is stable in air and possesses suitable mechanical and physicalproperties, like solubility in organic solvents.

Similarly, the conducting polymer polypyrrole exhibits conductivitiesranging from about 1 S/cm to about 100 S/cm. Investigators again foundthat placing substituent groups on either the nitrogen atom or a carbonatom of the heteroaromatic pyrrole ring decreases the conductivity ofpolypyrrole. For example, an unsubstituted polypyrrole, incorporatingthe tetrafluoroborate anion as the dopant compound, exhibits aconductivity of 40 S/cm, whereas the N-methyl derivative, incorporatingthe same dopant compound, exhibits a conductivity of 10⁻³ S/cm; thethree-methyl derivative of pyrrole exhibits a conductivity of 4 S/cm;the 3,4-dimethyl derivative, a conductivity of 10 S/cm; and the3,4-diphenyl derivative, a conductivity of 10⁻³ S/cm. Accordingly,substituted polypyrroles, although often demonstrating good physicalproperties, may not demonstrate a sufficient conductivity for use as theconducting polymer in a conductive analyte sensor.

As illustrated by the large conductivity drop in polypyrroles havingsubstituents positioned on the pyrrole ring, even substituents as smallas a methyl group introduce steric interactions sufficient toessentially destroy the conductivity of the polymer. However, as will bediscussed more fully hereinafter, if a conducting polymer has a suitablesubstituent group present on the monomer units, reproducible processingof the conducting polymer into a thin film of uniform thickness isfacilitated. Therefore, it would be desirable to provide a conductivesensor including a conducting polymer that is easy to process and thatalso exhibits a sufficiently high conductivity for sensitive andaccurate analyte determinations.

The inability to add a substituent to a polypyrrole withoutsignificantly reducing the conductivity of the polymer is critical froma perspective of polymer processing. Polypyrrole and presently-knownpolypyrrole derivatives are intractable polymers that are insoluble incommon organic solvents. Therefore, a polypyrrole cannot be processedconveniently. However, if a polypyrrole can be substituted withoutadversely affecting the electrical properties of the polymer, aprocessible polypyrrole may be developed.

The synthesis and conductivities of polypyrrole and substitutedpolypyrroles have been investigated extensively. The general referencescited below include the information discussed above and further generalinformation concerning polypyrroles. The representative referencesdiscussing the polypyrroles include:

M. S. Wrighton, Science 231, 32 (1986);

A. F. Diaz et al., J. Electroanal. Chem. 133, 233 (1982);

M. V. Rosenthal et al., J. Electroanal. Chem. and Interfac. Chem. 1, 297(1985);

G. Bidan et al., Synth Met. 15, 51 (1986);

E. M. Genies et al., Synth. Met. 10, 27 (1984/85); and

J. P. Travers et al., Mol. Cryst. Liq. Cryst. 118, 149 (1985).

Investigators also have found that steric interactions in substitutedpolythiophenes are somewhat less dominant than those observed insubstituted polypyrroles because the predominant destabilizinginteractions in substituted pyrroles involve the hydrogen atom of thepyrrole nitrogen. These steric interactions are absent in thepolythiophenes, and therefore, electronic effects are more predominantin substituted polythiophenes.

Polythiophene also is a well-studied, stable conducting polymer.Polythiophene resembles polypyrrole in that polythiophene can becyclized between its conducting, i.e. oxidized, state and itsnonconducting, i.e. neutral, state without significant chemicaldecomposition of the polymer and without appreciable degradation of thephysical properties of the polymer. Polythiophene, like polypyrrole,exhibits conductivity changes in response both to the amount of dopantcompound and to the specific dopant compound, such as molecular iodine,iridium chloride, arsenic(III) fluoride, phosphorus(V) fluoride,perchlorate, tetrafluoroborate, hexafluorophosphate, hydrogen sulfate,hexafluoroarsenate and trifluoromethylsulfonate.

Substituents placed on the heteroaromatic thiophene ring can affect theresulting conducting polymer. However, in contrast to pyrrole, ringsubstituents on thiophene do not seriously reduce the conductivity ofthe resulting heteroaromatic polymer. For example, it has been foundthat for 3-methylthiophene and 3,4-dimethylthiophene, the resultingsubstituted polythiophene exhibited an improved conductivity compared tothe unsubstituted parent polythiophene, presumably due to enhanced orderin the polymer chain of the substituted thiophene. Accordingly, unlikemany pyrroles, substituents can be included on the thiophene monomerunits to improve the processing properties of the resulting conductingpolymer without adversely affecting the conductivity of the conductingpolymer. Therefore, as a class, substituted polythiophenes arewell-suited for use as the conducting polymers in a conductive analytesensor.

The following are representative publications describing the synthesisand conductivity of polythiophene, substituted polythiophenes and therelated poly(thienylene vinylenes):

G. Tourillon, "Handbook of Conducting Polymers," T. A. Skotheim, ed.,Marcel Dekker, Inc., New York, 1986, p. 293;

R. J. Waltham et al., J. Phys. Chem. 87, 1459 (1983);

G. Tourillon et al., J. Polym. Sci. Polym. Phys. Ed. 22, 33 (1984);

G. Tourillon et al., J. Electroanal. Chem. 161, 51 (1984);

A. F. Diaz and J. Bargon, "Handbook of Conducting Polymers," T. A.Skotheim, ed., Marcel Dekker, Inc., New York 1986, p. 81;

G Tourillon et al., J. Phys. Chem. 87, 2289 (1983);

A. Czerwinski et al., J. Electrochem. Soc. 132, 2669 (1985);

K. Jen et al., J. Chem. Soc., Chem. Commun., 215 (1988); and

R. L. Elsenbaumer et al., Electronic Properties of Conjugated Polymers,400 (1987).

From the studies on the polyacetylenes, polypyrroles and polythiophenes,and from related studies on other conducting polymers, it becameapparent that a balance exists between the electronic effects and thesteric effects introduced by the substituent on the monomer unit thatrenders a polymer of a substituted five or six member heteroaromaticring more conducting or less conducting than the unsubstituted parentheteroaromatic compound. Therefore, it would be advantageous to utilizea conducting polymer having sufficient conductivity and suitableprocessing properties, such that the polymer can be used in a sensitiveconductive analyte sensor of a diagnostic device to accurately determinethe presence or concentration of a predetermined analyte in a liquidtest sample.

Accordingly, the present invention is directed to a conductive sensoruseful in a diagnostic assay for a predetermined analyte. Morespecifically, the conductive sensor is used in a diagnostic device todetermine the presence or concentration of a predetermined analyte, likeglucose, that is capable of interacting with an oxidase enzyme. Thepredetermined analyte and the oxidase interact in a reaction zone of theconductive sensor to generate, either directly or indirectly, a dopantcompound. The dopant compound then migrates to a detection zone of theconductive sensor to oxidatively dope a layer or film of conductingpolymer, thereby altering the conductivity of the layer of conductingpolymer. The change in conductivity of the conducting polymer then isdetected or measured by a microelectrode assembly, and can be correlatedto the concentration of the predetermined analyte in the test sample.

The prior art includes teachings related to diagnostic assays usingconducting polymers. However, the known prior art does not include anyreferences suggesting or anticipating the miniaturized conductive sensorof the present invention, or its method of use. Furthermore, althoughseveral references disclose the use of conducting organic polymers insensors, no known prior art reference discloses a sensor demonstratingthe sensitivity and accuracy of the conductive sensor of the presentinvention. The prior art sensors are based upon a direct interaction ofan analyte, usually a gas, with the conducting polymer. In contrast, thepresent conductive sensors utilize a response to a dopant compoundgenerated, either directly or indirectly, by an interaction between thepredetermined analyte and an oxidase enzyme. The dopant compound thenmigrates, or diffuses, into the layer of conducting polymer and altersthe conductivity of the layer of conducting polymer, thereby allowingthe determination of the presence or concentration of the predeterminedanalyte in the test sample.

The observed sensitivity of the conducting polymer to the concentrationof the dopant compound is important in the development of sensors thatare based upon oxidase enzymes. As will be demonstrated more fullyhereinafter, a very small amount of the dopant compound generates aresponse. Therefore, only a small fraction of the availablepredetermined analyte in the test sample, such as less than 1% of theavailable analyte, is interacted and converted into the dopant molecule.As a result, the problem of an oxygen limitation, inherent in usingoxidase enzymes, is overcome, and a sensitive determination of the totalanalyte concentration in the test sample is achieved.

The most common mode of interaction between an analyte and a conductingpolymer is to affect the state of oxidation of the organic conductingpolymer. As will be discussed more fully in the detailed description ofthe invention, the conductivity of the conducting polymer is related tothe degree of oxidation of the conducting polymer, with the degree ofoxidation, in turn related to the amount of dopant compound doping theconducting polymer. Therefore, the measured change in conductivity ofthe conducting polymer is correlated to the amount of dopant compoundoxidizing the conducting polymer. In accordance with the method of thepresent invention, the amount of dopant compound in the layer ofconducting polymer is directly related to the amount of predeterminedanalyte in the test sample. Consequently, a measurement of the rate ofconductivity change of the layer of conducting polymer also is ameasurement of the concentration of the predetermined analyte in thetest sample.

For example, M. K. Malmros et al., in the publication, "A SemiconductivePolymer Film Sensor for Glucose", Biosensors 3, pp. 71-87 (1987/88),suggest using polyacetylene in a biosensor to quantitatively detectglucose. Malmros et al. teach the conversion of glucose by glucoseoxidase to gluconic acid and hydrogen peroxide, and the subsequentconversion of iodide ion to molecular iodine, or triiodide anion, by theaction of lactoperoxidase. Malmros et al. further teach that the iodineso generated can be used to change the conductivity of a polyacetylenefilm, particularly as a modifier of the effect of the peroxide.

Malmros et al. however do not teach a biosensor that solves the problemsof assay interference and of oxygen limitation due to the low oxygenconcentrations in biological samples. Malmros et al. teach the physicaleffect of the enzymatically generated iodine doping the polyacetylenepolymer. The iodine is generated in solution through the addition ofsolution phase enzyme. Malmros et al. do not teach the incorporation ofthe enzymes or the iodide ion into solid phase films that can belaminated onto a conducting polymer film. Similarly, Malmros et al. donot teach a means of metering the test sample through a diffusionbarrier or measuring the early kinetics of the reaction.

Moreover, the method disclosed by Malmros et al. cannot be employed in abiosensor due to the inherent limitations of the polyacetylene film.Most importantly, a polyacetylene cannot be processed using eithersolution casting or low temperature melt processes. As a result, thinpolyacetylene films, such as 200 Å in thickness, cannot be produced. Athick film of conducting polymer greatly reduces the sensitivity of thefilm to a dopant compound by requiring more dopant compound to achievethe same level of doping. This in turn increases the amount ofpredetermined analyte in the sample that must be converted into iodineto achieve a response. As will be demonstrated more fully hereinafter,the ability to cast thin films of the conducting polymer is an importantaspect in providing reliable and accurate assay results.

Similarly, Malmros et al., in European Patent Application PublicationNo. 096,095, disclose an immunoassay utilizing doped conductingpolymers, wherein the resistance of the polymer varies in response tothe analyte in solution. Although Malmros et al. utilize a polyacetyleneconducting polymer in an immunosensor, Malmros et al. do not teach orsuggest assaying for a predetermined analyte capable of interacting withan oxidase enzyme to, either directly or indirectly, produce a dopantcompound such that the rate of the conductivity change of the conductingpolymer as the dopant compound oxidizes the conducting polymer iscorrelated to the concentration of the predetermined analyte in the testsample.

Wrighton et al., in European Patent Application Publication No. 185,941,disclose the use of conducting organic polymers as the active species ina chemical sensor. Wrighton et al. generally teach using the changes inphysical properties of the conducting polymer as the active transductioninto electrical signals. Specific examples cited in the patent includedetection of oxygen gas, hydrogen gas, pH and enzyme substrateconcentrations like glucose. The principal transduction mechanismdescribed by Wrighton et al. is the direct use of the change in polymerconductivity induced by oxidation or by reduction. In contrast to thepresent invention, Wrighton et al. include the reaction catalyst, likean enzyme, in the conducting polymer matrix. Accordingly, theinteraction occurs within the conducting polymer matrix. In the deviceand method of the present invention, the entire analyte-oxidaseinteraction occurs essentially in a reaction zone of the device togenerate the dopant compound; the dopant compound then migrates from thereaction zone to a detection zone that includes the layer of conductingpolymer. The rate of change of conductivity of the conducting polymerlayer as the dopant compound diffuses from the reaction zone to thedetection zone is used to measure the concentration of the predeterminedanalyte in the test sample. Wrighton et al. do not teach a method ofintegrating the glucose oxidation by oxygen into the oxidationproperties of the polymer, nor do Wrighton et al. teach the conversionof glucose to iodine through coupled enzymatic reactions. Furthermore,Wrighton et al. do not teach the use of solution processible polymers;all of the polymers used by Wrighton et al. are grown electrochemically.

Wrighton et al., in U.S. Pat. No. 4,717,673, disclose polymer-basedmicrosensors produced by anodic deposition of a conducting polymer ontoa gold or platinum electrode surface. Wrighton et al. in U.S. Pat. No.4,721,601 disclose microelectronic devices having electrodesfunctionalized with conducting polymers having specific properties. Thedevices disclosed in each patent measure resistance changes ofelectrochemically-grown polymers on electrode arrays having intergapspacings of less than two microns.

Elsenbaumer et al., in the publication "Processible, EnvironmentallyStable, Highly Conductive Forms of Polythiophene", Synth. Met., 18, pp.277-282 (1987), describe a series of conducting poly(3-alkylthiophene)polymers having sufficient conductivity, stability, mechanicalproperties and processibility for a variety of applications. Jen et al.,in the publication "Processible and Environmentally Stable ConductingPolymers", Polymeric Material, Vol. 13, pp. 79-84 (1985) also describepolythiophenes having good conductivity and mechanical properties.

Hotta et al., in "Novel Organosynthetic Routes to Polythiophene and ItsDerivatives", Synth. Met., 26, pp. 267-279 (1988) teach the synthesis ofpoly(3-alkylthiophene) polymers having long alkyl side chains. Suchconducting polymers are soluble in common organic solvents, areprocessible into uniform films and, when doped, exhibit excellentconductivity. Yoshino et al. in the three publications:

"Preparation and Properties of Conducting Heterocyclic Polymer Films byChemical Method", Jpn. Journ. Appl. Phys., 23, pp. 2899-2900 (1984);

"Electrical and Optical Properties of Poly(3-alkylthiophene) in LiquidState", Solid State Commun., 67, pp. 1119-1121 (1988); and

"Absorption and Emission Spectral Changes in a Poly(3-alkylthiophene)Solution with Solvent and Temperature", Jpn. Journ. Appl. Phys., 26, pp.L2046-L2048 (1987), describe the synthesis and properties of theconducting poly(3-alkylthiophene) polymers. Jen et al., in U.S. Pat. No.4,711,742, disclose doped and undoped conducting polymers that can besolubilized in organic solvents, with the resulting solution used toform conducting polymer films, including films of apoly(3-alkylthiophene).

Nagy et al., in the publication "Enzyme Electrode for Glucose Based onan Iodide Membrane Sensor," Analytica Chim. Acta., 66, pp. 443-455(1973), describe the detection of glucose by potentiometricallymonitoring the disappearance of iodide ion. Nagy et al. monitor thedecrease in iodide activity at the electrode surface, whereas thepresent invention monitors the amount of a predetermined analyte in thetest sample by measuring the rate of change of conductivity of theconducting polymer due to the generation of a dopant compound by anoxidase enzyme mediated interaction.

Mullen et al., in the publication, "Glucose Enzyme Electrode withExtended Linearity," Analytica. Chem. Acta., 183, pp. 59-66 (1986),describe a hydrogen peroxide-detecting electrode to assay whole bloodfor glucose. Mullen et al. disclose positioning a silane-treatedmembrane over a reactive enzyme layer to remove interferents and toextend the linearity of the electrode response to glucose concentrationin undiluted whole blood. Similarly, Vadgama, in European PatentApplication Publication No. 204,468, discloses a membrane for anenzyme-based electrode sensor to increase the range of linearity of thesensor response to generated hydrogen peroxide.

Other references relating to membranes, in general, or the detection ofglucose in a test sample in particular, include M. B. McDonell and P. M.Vadgama, "Membranes: Separation Principles and Sensing", SelectiveElectrode Rev., 11, pp. 17-67 (1989); L. C. Clark et al., "Long-livedImplanted Silastic Drum Glucose Sensors", Trans. Am. Soc. Artif. Intern.Organs, Vol. XXXIV, pp. 323-328 (1987); P. Vadgama et al., "The GlucoseEnzyme Electrode: Is Simple Peroxide Detection at a Needle SensorAcceptable?", in Implantable Glucose Sensors - The State of the Art,International Symposium, Reisensburg, pp. 20-22 (1987); P. Vadgama,"Diffusion Limited Enzyme Electrodes", Anal. Uses of ImmobilizedBiological Compounds for Detection, Medical and Industrial Uses, pp.359-377 (1988); and W. H. Mullen et al., "Design of Enzyme Electrodesfor Measurements in diluted Blood", Analytical Proceedings, 24, pp.147-148 (1987). These references also describe improved methods ofdetecting hydrogen peroxide in the assay of blood for glucose.

In addition to the above, the following references are representative ofthe state of the art of electromechanical sensors using heteroaromaticpolymers:

Y. Ikariyama et al., Anal. Chem., 58, 1803 (1986);

C. Nylander et al., Anal. Chem. Symp. Ser., 17, (Chem Sens) 159 (1983);

H. S. White et al., J. Am Chem. Soc., 106, 5317 (1984);

G. P. Kittlesen et al., J. Am. Chem. Soc., 106, 7389 (1984);

Malmros, U.S. Pat. No. 4,444,892, disclosing a device having an analytespecific binding substance immobilized onto a semiconductive polymer toallow detection of a specific analyte;

European Patent Application Publication No. 193,154, disclosingimmunosensors comprising a polypyrrole or polythiophene film having anantigen or antibody bound thereto; and

M. Umana and J. Waller, Anal. Chem., 58, 2979 (1986) disclose theocclusion, or trapping, of an enzyme, glucose oxidase, byelectropolymerizing pyrrole in the presence of the enzyme. Thepolypyrrole containing the occluded enzyme then can be used to detectglucose. However, the method of the present invention differs in twocritical respects. First, the detection mechanism in the presentinvention detects a generated oxidative dopant. The polymer filminitially is present in its reduced, nonconducting form, and becomesconductive only as the dopant compound is produced enzymatically. In thepublication of Umana and Waller, the polymer film initially isconductive and that conductivity is modulated by the enzymatic activityof the enzyme, that is serving as a dopant, and also by the generatedperoxide. The second critical difference is that, in the presentinvention, the oxidase enzyme is retained in a distinct reaction zonelayer that is in contact with the conducting polymer film in a detectionzone layer.

Investigators also have studied various other problems associated withelectrochemical sensors. For example, one major problem encountered inassays based on oxidase chemistry is the limited amount of molecularoxygen present in the system. In an oxidase catalyzed reaction, thepredetermined analyte reacts with an equimolar amount of molecularoxygen. When the supply of molecular oxygen is depleted, the reactionceases regardless of the presence of the oxidase enzyme and unreactedanalyte. If no further molecular oxygen can enter the system, anerroneously low assay for the predetermined analyte results. Ifmolecular oxygen can diffuse into the system, the oxidase-catalyzedreaction will continue, although slowly, until all the analyte isconsumed. In this case, an accurate analyte is achieved, but the timeneeded to achieve the accurate assay is impractically long.

Therefore, investigators have sought methods to overcome the problem ofoxygen limitation. One method is to mediate the oxidase-catalyzedreaction with a species other than oxygen. In this method, a compoundsuch as ferrocene, a ferrocene derivative, ferricyanide couples ortetrathiafulvalene/tetracyanoquinone is used as a replacement formolecular oxygen. These compounds perform in a manner similar tomolecular oxygen and are included in a sufficient amount such that allof the predetermined analyte in the test sample is oxidized. This methodwas used in an amperometric probe and is described by Cass et al. inAnal. Chem., 56, p. 607 (1984) and in Biosensors, Instrumentation andProcessing, The World Biotech. Report, Vol. 1, Part 3, p. 125 (1987).

Another disclosed method of avoiding the oxygen limitation problem is toeliminate oxygen and oxygen substitutes altogether, and allow a directelectron transfer from the enzyme to the electrode. This method isdisclosed by Y. Degani and A. Heller, J. Phys. Chem., 91, 1285 (1987).Furthermore, Vadgama, in European Patent Application Publication No.204,468, disclosed avoiding oxygen limitations by using a silane-treatedmembrane to restrict entry of the predetermined analyte into thereaction enzyme layer. In contrast, the device and method of the presentinvention avoid the oxygen limitation problem kinetically, that is bydetecting and measuring the concentration of the predetermined analytebefore oxygen limitations occur and by limiting the amount of thepredetermined analyte that contacts the reaction zone including theoxidase enzyme and the limited amount of molecular oxygen.

Another problem encountered in the design of a conductive sensor is theeffect of interfering compounds that often are present in a test sample,such as the presence of ascorbate ion in the assay of a biological fluidfor glucose. Investigators have found that in amperometric probes,interference from relatively easily oxidized compounds, such asascorbate ion, phenolics, uric acid, acetaminophen and salicylates,occurs because the interfering compound is oxidized at the anode.Investigators accordingly have attempted to eliminate the affect ofthese interfering compounds. A common technique is exemplified in thepublication of C. J. McNeil, et al., Biosensors, 3, p. 199-209(1987/88), wherein the electroactive species was functionalized to lowerits oxidation potential and thereby eliminate the interference in animmunosensor. I. Hannig et al., in "Improved Blood Compatibility at aGlucose Enzyme Electrode Used for Extra Corporeal Monitoring", Anal.Letters, 19(3&4), pp. 461-478 (1986), attempted to eliminate the effectsof interferents by utilizing a thick enzyme layer to convert arelatively large amount of the available glucose. The correspondinglarge response for glucose conversion essentially swamped theinterferent response.

In contrast, the method and device of the present invention utilizes aconductometric detection and measurement. Accordingly, an extremely lowvoltage can be used. The voltage is much lower than the oxidationpotential of the interfering compounds, and therefore the interferingcompounds are not oxidized. In addition, in the present invention, allchemical interactions occur in the reaction zone of the conductivesensor. The molecular iodine is generated in the reaction zone andmigrates to dope the conducting polymer in the detection zone.Therefore, no direct interference is possible at the electrode.

However, the generated molecular iodine dopant compound is capable ofinteracting with various serum components, like ascorbate. If asufficient amount of the molecular iodine interacts with serumcomponents rather than doping the polymer, interferences are observed.Therefore, the method of the present invention relies upon fast assaysto minimize interfering reactions of the generated molecular iodine.This is accomplished by the configuration of the sensor of the presentinvention, comprising a thin reaction zone and a thin detection zone,such that the molecular iodine is generated near the conducting polymerto quickly dope the conducting polymer before significant interferingreactions can occur. Furthermore, in the preferred embodiment of thepresent invention, it will be demonstrated that a semipermeable membraneutilized to meter the test sample into the reaction zone alsoselectively screens interfering compounds from the test sample, andtherefore precluding an interaction with the generated molecular iodine.

Therefore, the method of the present invention allows the accurate assayof a predetermined analyte that is responsive to oxidase chemistry. Themethod utilizes a diagnostic device that includes a conductive sensor,wherein the conductive sensor comprises a reaction zone and a detectionzone. The reaction zone of the conductive sensor is a thin filmincluding the reagents necessary to interact with the predeterminedanalyte and to generate a dopant compound. The detection zone includes afilm or layer of a conducting polymer and a microelectrode assembly suchthat the dopant compound can migrate to the detection zone to dope theconducting polymer, and such that the resulting change in conductivity,detected and measured by the microelectrode assembly, can be correlatedto the amount of predetermined analyte in the test sample. Theconductive sensor overcomes the disadvantages demonstrated by the priorart sensors, and therefore provides sensitive, accurate and reproducibleassays; provides a fast assay, such as within 30 seconds, and preferablywithin 10 seconds, from a small blood sample, such as from about 0.1 μlto about 5 μl; eliminates the oxygen limitation problem associated withoxidase chemistry; eliminates the problems associated with interferingcompounds present in the test sample; demonstrates excellent shelfstability; is economical and disposable; and is miniaturized and can bereproducibly manufactured by semiconductor processing techniques.

SUMMARY OF THE INVENTION

In brief, the present invention is directed to a diagnostic deviceincluding a conductive analyte sensor comprising a reaction zone and adetection zone, wherein the detection zone includes a conducting polymerand a microelectrode assembly. More particularly, the present inventionis directed to a conductive sensor that allows the sensitive andaccurate detection and measurement of a predetermined analyte in aliquid test sample, wherein the predetermined analyte is assayed by anoxidase interaction. In accordance with the method of the presentinvention, an interaction between the predetermined analyte and anoxidase enzyme occurs in the reaction zone of the conductive sensor toproduce, either directly or indirectly, a dopant compound that migratesto the detection zone of the sensor. The detection zone of the device isin laminar contact with the reaction zone and includes a layer or filmof conducting polymer that is oxidized by the dopant compound.Therefore, the conductivity of the conducting polymer layer is changed,and the change in conductivity of the conducting polymer layer isdetected and measured by the microelectrode assembly and is correlatedto the concentration of the predetermined analyte in the test sample.

The conductive sensors of the present invention utilize the uniqueelectrical properties of conducting polymers to determine the presenceand concentration of a predetermined analyte that is capable ofinteracting with a specific oxidase enzyme. In accordance with themethod and device of the present invention, the conductive sensorsinclude a layer of conducting polymer in a detection zone of the sensor.The conducting polymer is oxidized by a dopant compound generated,either directly or indirectly, as a result of an interaction between thepredetermined analyte and an oxidase enzyme in a reaction zone of thesensor. The dopant compound migrates from the reaction zone to thedetection zone of the device to oxidize the layer of conducting polymer.The conductivity of the polymer, therefore, is changed by theintroduction of the dopant compound into the conducting polymer layer,and the measurable conductivity change is detected and measured by amicroelectrode assembly in the detection zone and is correlated to theconcentration of the predetermined analyte in the test sample.

For example, hydrogen peroxide is produced in the reaction of glucoseoxidase with glucose in the presence of oxygen. Then, the hydrogenperoxide, in the presence of a peroxidase enzyme or a compound thatexhibits peroxidase activity, like a molybdenum(VI) transition metalcatalyst, can interact with iodide ions to generate molecular iodine.Each of these reactions is quantitative. Therefore, by measuring theamount of molecular iodine that is generated, the original amount ofglucose can be determined. In addition, the generated molecular iodineis a dopant compound, and the concentration of molecular iodine can bedetermined from the change in conductivity of a conducting polymer dopedby the molecular iodine. Accordingly, measuring the change, or the rateof change, in conductivity of the conducting polymer can be correlatedto the concentration of glucose in solution.

Consequently, it has been demonstrated that the conductive sensor of thepresent invention allows an accurate and sensitive electricaltransduction of an analyte-oxidase interaction, like a glucose-glucoseoxidase interaction. In accordance with an important feature of thepresent invention, the sensitive and accurate detection of apredetermined analyte, like glucose, results from the effect of theultimately generated dopant compound, like molecular iodine, upon theconductivity of a conducting polymer layer. This particular type ofreaction and detection method is known in the art. However, to date,utilizing this reaction in a conductive sensor has not provided asensitive and accurate assay for a predetermined analyte.

Therefore, and in accordance with the present invention, an interactionbetween a predetermined analyte and an oxidase enzyme, and a subsequentinteraction between the generated hydrogen peroxide, a dopant compoundprecursor and a peroxidase enzyme or a compound that exhibits peroxidaseactivity to generate a dopant compound, like molecular iodine, occurs ina reaction zone of a conductive sensor of a diagnostic device. Thedopant compound migrates from the reaction zone to contact andoxidatively dope a layer or film of conducting polymer present in adetection zone of the sensor. Consequently, the conductivity of theconducting polymer layer is changed, and the concentration of thepredetermined analyte is determined from the change in conductivity ofthe conducting polymer by a microelectrode assembly in the detectionlayer. Thus, to provide an accurate and sensitive assay by eliminatingoxygen limitation problems in the oxidase-based reaction, the change inconductivity of the conducting polymer layer in the conductive sensor ismeasured within about thirty seconds, and preferably within about 15seconds, after the test sample contacts the reaction zone of theconductive sensor. To achieve the full advantage of the presentinvention, the change in conductivity is measured from about 5 secondsto about 10 seconds after the test sample contacts the reaction zone ofthe sensor.

Therefore, it is an object of the present invention to provide a methodof determining the presence or concentration of a predetermined analytein a liquid test sample by utilizing a diagnostic device including aconductive sensor comprising a reaction zone in contact with a detectionzone. It also is an object of the present invention to provide a methodof determining the concentration of a predetermined analyte in a liquidtest sample wherein the predetermined analyte interacts with an oxidaseenzyme in the reaction zone of the sensor to generate, either directlyor indirectly, a dopant compound that oxidizes a layer or film ofconducting polymer in the detection zone of the sensor and that changesthe conductivity of the conducting polymer.

Another object of the present invention is to provide a method ofdetermining the concentration of a predetermined analyte in a testsample from the interaction of the analyte with an oxidase enzyme toproduce, either directly or indirectly, a dopant compound that oxidizesa conducting polymer, such that a detectable or measurable conductivitychange occurs in the conducting polymer and thereby establishes thepresence or concentration of the predetermined analyte in the testsample.

Another object of the present invention is to provide a method ofdetermining the presence or concentration of a predetermined analyte ina liquid sample comprising contacting a diagnostic test device with theliquid sample, wherein the diagnostic test device includes a conductivesensor comprising a reaction zone, wherein a portion of thepredetermined analyte interacts with an oxidase enzyme and otherreagents, if necessary, to generate a dopant compound, and a detectionzone in contact with the reagent zone, such that the dopant compoundmigrates to the detection zone to dope a layer of conducting polymerpresent in the detection zone and to cause a detectable or measurablechange in the conductivity of the conducting polymer; measuring thechange in conductivity of the layer of conducting polymer by amicroelectrode assembly present in the detection zone; and correlatingthe change in conductivity of the layer of conducting polymer to theconcentration of the predetermined analyte in the test sample.

Another object of the present invention is to provide a sensitive,miniaturized conductive sensor capable of assaying a sample volume inthe range of 0.1 μL to about 5 μL, and especially less than 1 μL, thataccurately senses the presence or concentration of a predeterminedanalyte in a test sample, wherein the predetermined analyte is capableof interacting with an oxidase enzyme; that essentially eliminates theeffects of interfering compounds present in the test sample; thateliminates the problem of oxygen limitation in the oxidase reaction; andthat provides an assay result within 10 seconds.

Another object of the present invention is to provide a sensitiveconductive sensor that accurately senses the presence or concentrationof a predetermined analyte in a liquid test sample, wherein thepredetermined analyte is capable of interacting with an oxidase enzyme,comprising: a) a reaction zone including a hydratable host matrixpermeable to the predetermined analyte and having homogeneouslyincorporated therein a suitable oxidase enzyme, a peroxidase enzyme or acompound that exhibits peroxidase activity, and any other necessaryreagents, like a dopant compound precursor, to generate a dopantcompound, and wherein the predetermined analyte interacts with theoxidase enzyme, peroxidase and other reagents, if present, to generate,either directly or indirectly, the dopant compound; b) a detection zonein contact with the reaction zone including a layer or film of aconducting polymer in contact with a microelectrode assembly, such thatthe dopant compound generated in the reaction zone can migrate to andoxidatively dope the film or layer of the conducting polymer; and c)means operatively connected to the microelectrode assembly of thedetection zone for measuring the conductivity of the conducting polymer.

Another object of the present invention is to provide a sensitiveconductive sensor that accurately senses the presence or concentrationof a predetermined analyte in a liquid test sample, wherein thepredetermined analyte is capable of interacting with an oxidase enzyme,comprising: a) semipermeable membrane capable of effectively separatingthe cellular material and interfering components from the test samplewhile allowing the predetermined analyte to diffuse through thesemipermeable membrane; b) a reaction zone including a hydratable hostmatrix permeable to the predetermined analyte and having homogeneouslyincorporated therein a suitable oxidase enzyme, a peroxidase enzyme or acompound capable of exhibiting peroxidase activity, and any othernecessary reagents, like a dopant compound precursor, to generate adopant compound, and wherein the predetermined analyte interacts withthe oxidase enzyme, the peroxidase enzyme or a compound that exhibitsperoxidase activity, and other reagents, if present, to generate, eitherdirectly or indirectly, the dopant compound; c) a detection zone incontact with the reaction zone including a layer or film of a conductingpolymer in contact with a microelectrode assembly, such that the dopantcompound generated in the reaction zone can migrate to and oxidativelydope the film or layer of the conducting polymer; and d) meansoperatively connected to the microelectrode assembly of the detectionzone for measuring the conductivity of the conducting polymer.

Another object of the present invention is to provide an accurate andsensitive miniaturized conductive sensor for determining the presence orconcentration of glucose in a liquid test sample of less than 1 μLcomprising a reaction zone wherein the hydratable host matrix is apolymer matrix, such as a gelatin matrix or a chitosan matrix,incorporating therein glucose oxidase, peroxidase and iodide ion tointeract with the glucose and to generate molecular iodine as the dopantcompound upon contact between the liquid test sample and the reactionzone.

Another object of the present invention is to provide an accurate andsensitive miniaturized conductive sensor for determining the presence orconcentration of a predetermined analyte capable of interacting with anoxidase enzyme comprising a detection zone comprising a layer, such as athin film, of a conducting polymer in contact with a microelectrodeassembly capable of measuring resistances as high as about 10⁹ ohms.

BRIEF DESCRIPTION OF THE DRAWINGS

The above and other objects and advantages and novel features of thepresent invention will become apparent from the following detaileddescription of the invention illustrated in the accompanying figuresdemonstrating the accurate and sensitive analyte assays achieved by theconductive sensor of the present invention wherein:

FIG. 1 is a partial side view in cross-section of a diagnostic device ofthe present invention comprising a capillary for introducing the testsample; and a conductive sensor comprising a reaction zone forinteracting the predetermined analyte with an oxidase enzyme and forgenerating the dopant compound, and a detection zone for detecting theamount of dopant compound generated in the reaction zone by measuringthe change in conductivity of a layer of conducting polymer present inthe detection zone with a microelectrode assembly;

FIG. 2 is a top view of a microelectrode assembly included in thedetection zone of the present invention showing the interdigitedmicroelectrodes and the spacing, or gap, between the interdigitedmicroelectrodes filled with the conducting polymer;

FIG. 3 is a partial side view in cross-section of a preferred embodimentof the conductive sensor of the present invention;

FIG. 4 is a dose response plot of current (amps) vs. time (seconds) forassays of standardized solutions including 30 mM (millimolar) potassiumiodide and either 25, 50, 100, 200, 300, 400 or 500 mg/dL (milligramsper deciliter) of glucose assayed by the method and conductive sensor ofthe present invention;

FIG. 5 is a plot of current (microamps at 10 sec.) vs. glucoseconcentration (mg/dL) showing the linear relationship between theconcentration of glucose in a standardized test sample and conductivityof the layer of conducting polymer exhibited by a conductive sensor ofthe present invention;

FIG. 6 is another dose response plot of current (amps) vs. time(seconds) for assays of standardized solutions including 30 mM potassiumiodide and either 25, 50, 100, 200, 300 or 500 mg/dL of glucose assayedby the method and conductive sensor of the present invention;

FIG. 7 is a plot of current (amps at 10 seconds) vs. glucoseconcentration (mg/dL) showing the linear relationship between theconcentration of glucose in a standardized test sample and conductivityof the layer of conducting polymer exhibited by a conductive sensor ofthe present invention;

FIG. 8 is a dose response plot of current (amps) vs. time (seconds) forassays of standardized solutions including 150 mM potassium iodide andeither 5, 90, 150, 300 or 500 mg/dL of glucose assayed by the method andpreferred embodiment of the conductive sensor of the present invention;

FIG. 9 is a plot of current (microamps at 10 sec.) vs. glucoseconcentration (mg/dL) showing the linear relationship between theconcentration of glucose in a standardized test sample and conductivityof the layer of conducting polymer exhibited by a conductive sensor ofthe present invention; and

FIG. 10 is a dose response plot of current (amps) vs. time (seconds) forassays of standardized solutions including 100 mM tetraethylammoniumiodide and either 100, 200 or 400 mg/dL of glucose assayed by the methodand conductive sensor of the present invention.

DETAILED DESCRIPTION OF THE INVENTION

In accordance with the method and device of the present invention, adiagnostic device includes a conductive sensor comprising a reactionzone and a detection zone, wherein the detection zone includes a layerof a conducting polymer and a microelectrode assembly, to determine thepresence or concentration of a predetermined analyte in a liquid testsample. Although conductive sensors have been studied extensively, theuse of a conductive sensor in a diagnostic device to assay for apredetermined analyte, like glucose, has been impeded by severalproblems, including poor conducting polymer properties, impracticalmethods of manufacturing the conductive sensor, the inability toaccurately test for a predetermined analyte, insensitivity to lowconcentrations of the predetermined analyte, irreproducible analyteassays, interferences associated with compounds often found in testsamples, and oxygen limitation problems in assays based on oxidasechemistry. As will be described more fully hereinafter, the method anddevice of the present invention surprisingly and unexpectedly overcomemany of the problems previous investigators encountered in attempts toincorporate a miniaturized conductive sensor into a diagnostic device.

In general, the conductive sensor of the present invention allows thedetection and measurement, or monitoring, of a predetermined analyte ina liquid test sample. More specifically, an analyte capable of detectionor measurement, either directly or indirectly, by an oxidase chemistrycan be assayed by the conductive sensor of the present invention. Forexample, glucose, cholesterol, alcohol and other analytes capable ofinteracting with a suitable oxidase enzyme can be assayed by the methodand device of the present invention. In addition, immunological analytescan be measured in competitive, displacement and sandwich ELISA formatsby labelling with an appropriate oxidase enzyme. The method and deviceof the present invention are especially useful in the in vitro detectionand measurement of glucose in a test sample. Accordingly, and as will bedemonstrated more fully hereinafter, the method and device of thepresent invention provide more simple, more accurate and more sensitiveanalyte assays than prior art methods and devices utilizing conductivesensors.

Accordingly, the method and device of the present invention are basedupon the oxidative doping of a conducting polymer, like apoly(3-alkylthiophene), present in the detection zone of the conductivesensor, by a dopant compound, such as molecular iodine, that isgenerated in the reaction zone of the sensor. As will be demonstratedmore fully hereinafter, a poly(3-alkylthiophene), as depicted instructural formulas (I) (undoped) and (II) (doped), is a particularlyuseful conducting polymer in the method and device of the presentinvention. However, it should be understood that other conductingpolymers demonstrating sufficient conductivity, i.e. detectable ormeasureable, and having sufficient stability, mechanical properties andprocessibility also are useful in the device and method of the presentinvention. For example, another class of conducting polymers exhibitingproperties useful in the device and method of the present inventionincludes the poly(thienylene vinylene) polymers, depicted in generalstructural formula (III), and the related poly(furylene vinylene)polymers. A poly(thienylene vinylene) film doped with molecular iodineexhibited a conductivity of 62 S/cm, and a poly(furylene vinylene) filmdoped with molecular iodine exhibited a conductivity of 36 S/cm.##STR1##

As previously discussed, the oxidized form of a conducting polymer, i.e.structure (II), demonstrates a dramatic increase in electricalconductivity compared to the reduced form of the conducting polymer,i.e. structure (I). For example, a conducting polymer in its completelyreduced form is insulating, and exhibits a conductivity of about 10⁻⁷S/cm. However, if the conducting polymer is fully oxidized, theconductivity can increase to as high as about 10³ S/cm, depending uponthe chemical structure of the conducting polymer and the morphologicalcondition of the conducting polymer layer. Furthermore, as disclosed byBurks and Hodge, in J. Chem. Phys., 83, p. 5796 (1985), a conductingpolymer requires approximately 10⁻¹¹ moles of dopant compound per squarecentimeter of surface area of conducting polymer to demonstrate an orderof magnitude change in the conductivity of a 200 Å (Angstrom) thick filmof conducting polymer. Accordingly, this very sensitive conductivityresponse of the conducting polymer, especially a thin film of conductingpolymer, to the concentration of the dopant compound doping theconducting polymer layer allows a conductive sensor of the presentinvention to provide sensitive and accurate assays for a predeterminedanalyte.

The increased conductivity of a conducting polymer upon oxidative dopingby a dopant compound, such as molecular iodine or arsenic triflouride,is a well recognized electrical property of a conducting polymer.However, it has been difficult to provide a uniform, thin film of aconducting polymer. If a thin film of polymer is available, less dopantcompound is needed to provide a detectable or measurable increase inconductivity. Surprisingly and unexpectedly, the conductive sensor andthe method of the present invention provide a uniform, thin film ofconducting polymer, such as from about 100 Å to about 1500 Å inthickness. Therefore, the sensitivity of the assay is increased becauseless dopant compound must be generated for a detectable response; andthe accuracy of the assay is increased because the electrical responseof a layer of the conducting polymer is increased, therefore makingdetection measurements easier and more reliable. In addition, and aswill be demonstrated more fully hereinafter, the ability to provide athin, uniform layer of a conducting polymer helps eliminate the oxygenlimitation problem found in oxidase-based assays. The sensitivedetection provided by a thin film of conducting polymer allows a methodwherein only a very minor portion of the analyte in the test sampleinteracts with the enzyme and the available oxygen. Without thesensitivity provided by a thin film of conducting polymer, the change inconductivity provided by such a small conversion of analyte could goundetected.

An embodiment of a conductive sensor of the present invention isillustrated in FIG. 1, wherein a diagnostic test device 10 utilizes aconductive sensor to assay a liquid test sample for a predeterminedanalyte in general, and for glucose in particular. In FIG. 1a, a liquidtest sample, such as a biological fluid that includes glucose, likewhole blood, blood serum, blood plasma, lacrymal fluid, interstitialfluid, or urine, is introduced into the diagnostic test device 10 by acapillary tube 12. Blood is drawn by capillary action through thecapillary tube 12 in the direction of the arrow in a small noninvasiveamount, such as from about 0.1 μL (microliters) to about 5 μL, from asmall, fresh wound. The test sample then contacts a surface of areaction zone 14 of the conductive sensor that is in contact withcapillary tube 12.

The reaction zone 14 comprises a hydratable host matrix, such as agelatin matrix or a chitosan matrix, that is from about 0.1μ (micron) toabout 10μ, and preferably from about 0.2μ to about 5μ, in thickness; andthat is capable of being hydrated quickly, such as within from about 1second to about 5 seconds, by the test sample. If the test sample iswhole blood, the hydratable host matrix of reaction zone 14 preferablyscreens, or precipitates, the cellular material, including the red bloodcells, from the serum or plasma, and fixes the cellular material in thehydratable host matrix of reaction zone 14.

In FIG. 1b, the reaction zone 14 is rapidly hydrated upon contact withthe test sample. The rapid hydration of the hydratable host matrixproduces a thin, continuous film 18 from about 2μ to about 10μ inthickness. The glucose concentration in the thin, continuous film 18 isessentially identical to the glucose concentration in the test sample.Therefore, an important feature of the present invention is to provide ahydratable host matrix that hydrates sufficiently fast to insure thatthe reaction zone 14 and the test sample have essentially the sameconcentration of predetermined analyte.

The chemical reagents necessary to interact with the predeterminedanalyte of interest and to form the dopant compound are incorporatedinto the hydratable host matrix of reaction zone 14. Accordingly, in theassay for glucose, the hydratable host matrix of reaction zone 14 hasincorporated therein assay reagents including glucose oxidase,peroxidase or a compound capable of exhibiting peroxidase activity, andiodide ions such that the following chemical interactions (1) and (2)can occur within reaction zone 14 after the test sample contacts andquickly hydrates the substrate matrix. ##STR2##

In the embodiment of the present invention illustrated in FIG. 1, instep (1) of the above interaction sequence, a portion of the glucose inthe test sample interacts in reaction zone 14 with molecular oxygen (O₂)through the mediation of the catalyzing enzyme, glucose oxidase, togenerate hydrogen peroxide. It should be understood that the hydratablehost matrix includes only a limited amount of molecular oxygen. Thisamount of oxygen is insufficient to convert all of the glucose in thetest sample to hydrogen peroxide. Therefore, in accordance with animportant feature of the method of the present invention only a smallamount of the glucose in the test sample is reacted, and the assaydetection is performed before the amount of molecular oxygen in thehydratable host matrix is depleted.

The hydrogen peroxide (H₂ O₂) generated in step (1), then interacts withthe iodide ion (I⁻) present in the hydratable host matrix through themediation of the peroxidase enzyme to form molecular iodine (I₂) in step(2) of the chemical interaction. The molecular iodine then serves as thedopant compound. In the embodiment illustrated in FIG. 1, only a smallamount of the total available glucose present in the test sample, suchas less than about 1%, and as low as about 0.1%, of the total glucosepresent in the test sample, is converted by the enzyme interactions ofsteps (1) and (2) to produce molecular iodine within reaction zone 14 oftest device 10. Surprisingly, the conductive sensor is sufficientlysensitive to detect the conductivity change provided by this smallamount of generated molecular iodine, and the conductive sensor is soconfigured such that the detection measurement can be made before theoxygen concentration in the reaction zone 14 is depleted.

Therefore, in accordance with the device and method of the presentinvention, a sufficient amount of ambient oxygen is present in thereaction zone 14 to allow a sufficient amount of the glucose in the testsample to interact with the oxidase enzyme and generate a sufficientamount of molecular iodine to dope a layer or film of conducting polymer16 present in a detection zone of the conductive sensor of test device10 and cause a detectable or measurable conductivity change in the layeror film of conducting polymer 16 (FIG. 1c). Surprisingly andunexpectedly, the method and device of the present invention provide aconductive sensor that is sufficiently sensitive and operatessufficiently quickly such that low concentrations of generated moleculariodine are detected. Therefore, it is not necessary to convert theentire amount of glucose in the test sample to molecular iodine.Accordingly, the sensitivity of the conductive sensor, in addition to aconfiguration that allows a conductivity measurement before the oxygensupply is depleted, overcomes the oxygen limitation found in prior artmethods and devices.

As previously stated, the production of molecular iodine or a similardopant compound requires an oxygen-based chemistry. Furthermore, becauseof the limited amount of ambient oxygen present in the hydratable hostmatrix of reaction zone 14, only a small amount of the glucose presentin the test sample is reacted enzymatically to eventually generatemolecular iodine. The method of the present invention, therefore, isbased upon a kinetic measurement of molecular iodine formation, whereinthe initial rate of glucose interaction, or equivalently moleculariodine production, is measured prior to a significant depletion of thesupply of ambient oxygen. The length of time prior to a significantdepletion of ambient oxygen is related to the amount of glucose oxidaseincorporated into reaction zone 14, the thickness of both the reactionzone 14 and the layer of conducting polymer 16 in the detection zone,the availability of ambient oxygen and the glucose concentration of thetest sample. However, in considering all of these parameters, it hasbeen found that from about ten seconds to about thirty seconds is thetypical length of time before the ambient oxygen supply is significantlydepleted.

The molecular iodine produced in the reaction zone 14 migrates to adetection zone of the conductive sensor that is in contact with reactionzone 14 (FIG. 1c). The detection zone includes a film or layer ofconducting polymer 16 and a microelectrode assembly 20. The top surfaceof the film of conducting polymer 16 is in contact with the reactionzone 14, and the bottom surface of the conducting polymer 16 is incontact with the microelectrode assembly 20. In general, in thedetection zone, the molecular iodine oxidatively dopes the film or layerof conducting polymer 16 and the conductivity of the film or layer ofconducting polymer 16 increases. A conductometric measurement is madeunder a constant applied potential between two microelectrodes presentin the microelectrode assembly 20 of the detection zone, and theincrease, or the rate of increase, of the conductivity of the film orlayer conducting polymer 16 is correlated to the glucose concentrationof the test sample. To achieve the full advantage of the presentinvention, the rate of increase in conductivity is measured before thesupply of ambient oxygen is significantly depleted, such as from within5 seconds to 30 seconds, and preferably from within 5 seconds to 10seconds, after the test sample contacts the reaction zone 14 of theconductive sensor of test device 10.

The sensitivity and accuracy of the device and method of the presentinvention are directly related to the physical and chemical propertiesof the components comprising the reaction zone 14 and comprising thedetection zone of the test device 10. For example, in regard to thedetection zone, the layer of conducting polymer 16 that is doped by themolecular iodine is included in the detection zone as a film or layerpositioned in laminar contact with the microelectrode assembly 20. Inparticular, the microelectrode assembly 20 includes an interdigited pairof metal electrodes with an insulating spacing of from about 5μ to about300μ, and preferably from about 5μ to about 250μ. To achieve the fulladvantage of the present invention, the interdigited pair of metalelectrodes has as small an insulating spacing as possible, such as fromabout 5μ to about 15μ. The microelectrode assembly 20 can be anyelectrode assembly capable of measuring a conductivity in the range offrom about 10⁻⁷ S/cm to about 10⁻² S/cm of a film or layer of conductingpolymer 16 approximately 100 Å to approximately 1500 Å in thickness.Although a suitable microelectrode assembly 20 can be any one of avariety of configurations, an especially suitable configuration isillustrated in FIG. 2, and includes a pair of interdigited electrodeshaving gap spacings ranging from about 5μ to about 15μ, wherein the morenarrow the gap spacings, the more sensitive the conductivitymeasurement.

In particular, the microelectrode assembly 30 illustrated in FIG. 2comprises a base 32 of a smooth, nonconductive material, like siliconmetal, ceramic or glass. The base 32 has interdigited patterns of aconductive material 34 and 36, like a metal, applied to the top surfaceof the base 32. The interdigited patterns of conductive material 34 and36 are conductively connected to conductive contacting pads (not shown)on the bottom surface of the base 32 by conductive vias (not shown)incorporated into the base 32. The interdigited patterns of conductivematerial 34 and 36 serve as the microelectrodes of the microelectrodeassembly of the conductive sensor. Therefore, when a layer or film ofconducting polymer is applied to the microelectrode assembly 30, thefilm or layer of conducting polymer bridges, or fills, the gap 38between the interdigited patterns of conductive material 34 and 36 and achange in conductivity of the layer or film of conducting polymer in thegap 38 is detected by the microelectrodes comprising the interdigitedpatterns of conductive material 34 and 36. The manufacture of amicroelectrode assembly 30 will be discussed more fully hereinafter.

In further regard to FIG. 1, in addition to the microelectrode assembly20, another essential component in the detection zone of the conductivesensor is a thin, uniform layer or film of conducting polymer 16. Theconducting polymer included in the layer or film of conducting polymer16 usually is a heteroaromatic conducting polymer, like a polypyrrole, apoly(thienylene vinylene), a poly(furylene vinylene), a polyfuran or apolythiophene. However, a carbocyclic aromatic conducting polymer, likepolyaniline, also is envisioned as useful in the method and device ofthe present invention. As will be discussed more fully hereinafter, asubstituted polythiophene is the preferred conducting polymer becausethe substituted polythiophenes possess sufficient electrical propertiesand physical properties, like solubility in organic solvents, for asensitive and accurate detection of the dopant compound and for easy anduniform manufacture of a thin film of conducting polymer.

Polythiophenes and polypyrroles containing long, flexible hydrocarbonchains at the 3-position of the heterocyclic rings have demonstratedsolution processibility. These 3-position substituents provideprocessibility without significantly impairing charge transport in thedoped polymers. Even with rather large substituents, the conductivitiesof the doped polymers can be high. As previously stated, in addition tothe chemical identity of the conducting polymer, the thickness of thelayer or film of conducting polymer 16 applied onto the microelectrodeassembly 20 is important because the thickness of the layer or film ofconducting polymer 16 is directly related to the sensitivity of theresponse of test device 10 to the amount of predetermined analyte in thetest sample. In turn, the ability to provide a thin film of conductingpolymer is directly related to the processibility of the polymer.

As the thickness of the layer or film of conducting polymer 16decreases, the less molecular iodine is required to oxidatively dope thelayer or film conducting polymer 16 to a particular conductivity value.As a result, it is desirable to have as thin a layer or film ofconducting polymer 16 as possible, while retaining suitablemorphological properties. Accordingly, a layer or film of conductingpolymer 16 having a thickness of from about 100 Å to about 2000 Å hasbeen found useful in the device and method of the present invention.However, a layer or film of conducting polymer 16 having a thickness ofup to about 10,000 Å can be used in the method and device of the presentinvention, although the sensitivity of the detection of molecular iodineand the accuracy of the assay may be decreased. Preferably, the film orlayer of conducting polymer 16 has a thickness of from about 100 Å toabout 1500 Å. However, a lesser or a greater thickness of the layer orfilm of conducting polymer 16 can be utilized as long as the gap betweenthe interdigited patterns of conductive material in the microelectrodeassembly 20 is bridged, or filled, with the conducting polymer and aslong as the interdigited patterns of conductive material are completelycovered by the film or layer of conducting polymer 16. If the film orlayer of conducting polymer 16 is too thick, a large test sample, moreassay reagents and a longer assay time may be needed to achieve asensitive and accurate assay.

The choice of conducting polymer used in the film or layer of conductingpolymer 16 also is important because the conducting polymer preferablyis easily processible when in solution, and exhibits rapid, facileoxidative doping upon exposure to a suitable dopant compound. Aparticularly useful class of polymers is the poly(3-alkylthiophene)polymers. Polyalkylthiophenes including an alkyl substituent with atleast four carbon atoms exhibit significant solubility in many organicsolvents including chloroform, methylene chloride, xylene andtetrahydrofuran. This solubility in organic solvents permits the use ofvarious processing techniques that reliably and uniformly deposit thinpolymer layers, or films, onto a substrate. These processing techniquesinclude batch process techniques like spin coating, film casting, inkjet printing and similar batch process techniques. In particular, spincoating is a preferred method of depositing a layer or film of theconducting polymer onto a substrate electronic template. Similarly,polyaniline compounds can be solubilized in concentrated sulfuric acidand cast as a thin uniform film; and poly(thienylene vinylene) andpoly(furylene vinylene) prepolymers are soluble and processible in thepre-polymer stage and can be cast, then polymerized. Although each ofthese types of polymers can be processed into thin films, the processingis relatively difficult. Accordingly, polymers such as thepolyalkylthiophenes that are soluble in common organic solvents forsimple casting into a film are preferred. In addition, as will bedemonstrated more fully hereinafter, the reagents and the hydratablehost matrix comprising the reaction zone 14 can be formed into auniform, thin layer by the same batch-type processes used to form a thinlayer of conducting polymer.

In addition to suitable processing properties, the electrical propertiesof the poly(3-alkylthiophene) polymers, like poly(3-octylthiophene),also are suitable for use in the method and device of the presentinvention. Although other conducting polymers possess equal or superiorelectrical properties, poly(3-alkylthiophene) polymers are preferredpolymers because of their suitable electrical properties and theirsolubility in organic solvents. For example, a thin film of a3-alkylthiophene polymer dopes very quickly upon exposure to a dopantcompound, like molecular iodine. The films of poly(3-alkylthiophene)polymers wherein the alkyl group includes from about four to abouttwenty carbon atoms generally are doped more quickly than films ofpoly(3-alkylthiophene) polymers wherein the alkyl group includes lessthan about four carbon atoms, such as poly(3-methylthiophene).Therefore, to achieve an optimum doping of the layer of conductingwithin a short time period of less than about 30 seconds, it ispreferred that the alkyl chain on the poly(3-alkylthiophene) includesfrom about six to about twelve carbon atoms. To achieve the fulladvantage of the present invention, the alkyl group of thepoly(3-alkylthiophene) includes from about six to about nine carbonatoms. Furthermore, a comonomer, such as thiophene, 3-methylthiophene or3-ethylthiophene, can be included in the conducting polymer as a meansof optimizing the electrical properties and the processing properties ofthe conducting polymer.

The morphological and electrical properties of the film or layer ofconducting polymer included in the conductive sensor also can bemodified by incorporating a surfactant or an inert, nonconductingpolymer, such as, but not limited to, polyethylmethacrylate,polyacrylonitrile, polyethylene oxide, polyvinyldene chloride, nylon,polystyrene, polyacrylic acid, polyacrylamide, polyester, and similarnonconducting polymers, into the layer or film of conducting polymer.For example, an inert, nonconducting polymer, such aspolymethylmethacrylate (PMMA), can be solubilized in an organic solvent,like chloroform, with a conducting polymer to form a casting solution.If the conducting polymer is included in the casting solution in aconcentration sufficient to ensure electrical percolation, i.e. at leastabout 25% poly(3-octylthiophene) for a poly(3-octylthiophene)/PMMAcasting solution, the final conductivity of the film of conductingpolymer after oxidative doping by molecular iodine is approximatelyequal to the conductivity demonstrated by a doped 100% 3-alkylthiophenepolymer film.

Alternatively, a copolymer of 3-octylthiophene and 3-methylthiophene canbe used as the conducting polymer. The poly(3-alkylthiophene) copolymerincorporates the excellent electrical properties and affinity formolecular iodine of poly(3-methylthiophene) and the excellent solubilityand processing properties of poly(3-octylthiophene) to provide aparticularly useful conducting polymer. It has been found thatcopolymerizing a mixture including at least 90% by weight3-octylthiophene and up to 10% by weight of 3-methylthiophene provides acopolymer that exhibits sufficient solubility for easy processing andexcellent doping and electrical properties. For example,poly(3-octylthiophene) doped with ferric chloride demonstrated aconductivity of 1 S/cm, whereas, the copolymer (50:50)poly(3-octylthiophene-co-3-methylthiophene) doped with ferric chlorideexhibited a conductivity of 20 S/cm. However, copolymers including sucha high amount of 3-methylthiophene are not easily processed. Copolymersof various 3-alkylthiophene are described in Jen et al., U.S. Pat. No.4,711,742.

Therefore, in general, the selection of a particular conducting polymeris limited only in that it exhibits sufficient conductivity uponoxidative doping; is sufficiently soluble in organic solvents; and isprocessible into a uniform layer. In general, it has been found that anyconducting polymer that exhibits a solubility of at least 2 mg/ml(milligrams per milliliter) of solvent, either aqueous or organic, canbe used to cast the layer of film or conducting polymer. Preferably, theconducting polymer exhibits a solubility of at least 5 mg/ml of solvent.It also is advantageous to select a conducting polymer that, whencombined with a nonconducting polymer, provides a film or layer ofconducting polymer that demonstrates essentially the same conductivityas a film or layer of the conducting polymer alone. Accordingly, by ajudicious selection of a poly(3-alkylthiophene) and a nonconductingpolymer, or by a judicious selection of a poly(3-alkylthiophene)copolymer, a film or layer of conducting polymer that is easilyprocessed, that exhibits a high conductivity and that exhibits afavorable doping-undoping equilibrium can be provided.

In addition to the particular chemical and physical properties requiredof the film or layer of conducting polymer 16 and the microelectrodeassembly 20 in the detection zone of the conductive sensor of the testdevice 10 in FIG. 1, the reaction zone 14 also should possess suitablechemical and physical properties for the device and method of thepresent invention to detect and accurately measure the predeterminedanalyte in the test sample. In general, the reaction zone 14 of theconductive sensor of test device 10 illustrated in FIG. 1 is a layerfrom about 0.1μ to about 10μ, and preferably from about 0.2μ to about5μ, in thickness, when dry. To achieve the full advantage of the presentinvention, the reaction zone is a layer from about 0.2μ to about 3μ inthickness, when dry. However, the thickness of the layer of hydratablehost matrix is limited only in that the layer is sufficiently thick toincorporate the necessary amounts of the oxidase enzyme, peroxidase anddopant compound precursor; and is sufficiently thin such that the assaycan be performed within 30 seconds without interference from othercompounds often found in a test sample. If the hydratable host matrix istoo thick, the molecular iodine is generated relatively far from thelayer of conducting polymer, therefore requiring more time for themolecular iodine to migrate to the layer of conducting polymer andproviding time for the generated molecular iodine to interact withinterfering compounds in the test sample, like ascorbate ion, therebyproviding an inaccurate assay.

The reaction zone 14 comprises a hydratable host matrix that uniformlyincorporates the oxidase enzyme, the peroxidase enzyme or a compoundthat exhibits peroxidase activity, like a molybdenum(VI) transitionmetal catalyst, and any necessary dopant compound precursors, likeiodide ion. A suitable hydratable host matrix includes materials such asgelatin, silk fibroins, chitosan, collagen, and polyacrylamide; orcombinations thereof. The preferred hydratable host matrix is gelatin,chitosan or silk fibroins, or a combination thereof. To achieve the fulladvantage of the present invention, gelatin or chitosan is included inthe hydratable host matrix.

In accordance with an important feature of the present invention, thedry, hydratable host matrix is capable of rapid hydration upon contactwith the liquid test sample, such as within about 5 seconds. Inparticular, the method of the present invention utilizes an earlyconductivity measurement, i.e. within about 30 seconds, and preferablywithin about 10 seconds, after the test sample contacts the reactionzone 14, to determine the rate of molecular iodine formation, orequivalently, the rate of glucose conversion. Therefore, it is importantthat the dry, hydratable host matrix hydrates before a substantialinteraction between the predetermined analyte and the oxidase enzymeoccurs. It also is important that the diffusion rates of allinteractants and interaction products, like molecular iodine, throughthe hydratable host matrix are sufficiently high such that theinteractions can proceed quickly, and that the generated moleculariodine can quickly and effectively migrate to the detection zone ofconductive sensor to dope the layer or film of conducting polymer film16.

In addition, because the oxidase enzyme is not present in the reactionzone 14 in an excess amount, it also is important that the oxidaseenzyme remains stable within the hydratable host matrix over potentiallylong storage periods. Finally, the processing technique used to applythe reaction zone 14 to the conductive sensor must not disrupt theintegrity of the film or layer of conducting layer 16 present in theunderlying detection zone. It also is envisioned that the oxidaseenzyme, peroxidase enzyme and iodide ion, or other dopant compoundprecursor, can be isolated from one another within the dry hydratablehost matrix, as long as these components diffuse sufficiently rapidlysuch that the host matrix is homogeneous when it is hydrated by theliquid test sample.

The method of the present invention, used to assay a test sample for apredetermined analyte capable of undergoing an interaction with anoxidase enzyme, is both new in the art and is distinct from the otherknown electrochemical sensors for such analytes. For example, thecommercial prior art sensors include amperometric glucose probes andother investigators are attempting to miniaturize potentiometric glucoseprobes. In addition to differing from the prior art devices and methods,the present invention demonstrates advantages that help overcome themajor problems and disadvantages common to most, if not all, of theprior art electrochemical glucose and related sensors.

For example, a major problem in electrochemical glucose and relatedsensors is the need for a sufficient amount of oxygen to interact withthe entire amount of analyte present in the test sample. Originally, theoxygen limitation problem was overcome by diluting the blood or serumsample. Although, early electrochemical sensors depended upon dilutionto eliminate oxygen limitations, dilution limited the applicability ofelectrochemical sensors in decentralized test markets. Accordingly, toeliminate oxygen limitations and avoid test sample dilutions,investigators began incorporating a mediator other than oxygen for theelectron transfer reactions of glucose oxidase. Such oxygen substitutesinclude 1,1'-dimethylferrocene and ferricinium ion mediators. Otherinvestigators relied upon direct electron relays between the oxidaseenzyme and the electrode surface.

However, the device and method of the present invention eliminate theproblem of oxygen limitation more easily and provide fast assaysperformed on small, undiluted test samples. In the embodimentillustrated in FIG. 1, the improved sensitivity of the device and methodresults from a technique that utilizes a kinetic measurement of the ratethe initial analyte-oxidase interaction before an appreciable amount ofthe ambient oxygen supply, or an appreciable amount of the analyte inthe test sample, has been consumed in the interaction. In general, onlyfrom about 0.1% to about 1% of the available predetermined analyte inthe test sample has been consumed at the time the conductometricmeasurement is made. Therefore, the oxygen limitation problem isovercome. Accordingly, the device and method of the present invention donot require oxygen-substitute mediators or additional manipulativesteps, like dilution. In addition, the analyte assay is fast, simple,sensitive and accurate.

Accordingly, the method of the present invention comprises introducing atest sample, such as from about 0.1 μL to about 5 μL, and usually lessthan 1 μL, of a whole blood sample, into the capillary tube 12 of testdevice 10 illustrated in FIG. 1; then determining the change ofconductivity in the layer of conducting polymer 16 in the detectionlayer within a time period of from about 5 sec. to about 30 sec. afterthe test sample contacts the reaction zone 14 of the conductive sensorof the test device 10. The change in conductivity of the layer or filmof conducting polymer 16 is measured by the microelectrode assembly 20and can be correlated to the amount of predetermined analyte in the testsample by comparison to the change in conductivity of a layer or film ofconducting polymer 16 exhibited by standardized solutions of thepredetermined analyte. Accordingly, a fast, simple, sensitive andaccurate assay for glucose, or other analytes capable of interactingwith an oxidase enzyme, is provided.

The preferred embodiment of the present invention is illustrated in FIG.3, wherein a conductive sensor 40 is utilized to assay for apredetermined analyte, such as glucose, in a test sample 48. Theconductive sensor 40, like the conductive sensor in the test device 10illustrated in FIG. 1, includes a reaction zone 44 in laminar contactwith a detection zone including a film or layer of a conducting polymer46 and a microelectrode assembly 50 that are essentially identical tothe reaction zone 14 and the detection zone of the conductive sensor oftest device 10 of FIG. 1 described above. For example, reaction zone 44includes the necessary oxidase enzyme to interact with the predeterminedanalyte of interest, peroxidase and a dopant compound precursor, ifnecessary, to generate a dopant compound. However, unlike the conductivesensor of test device 10 in FIG. 1, the reaction zone 44 of conductivesensor 40 includes an excess amount of oxidase enzyme and peroxidase toconvert a sufficient amount of the predetermined analyte in the testsample 48 to the dopant compound. In addition, the detection zoneincludes a layer or a film of conducting polymer 46 that is oxidativelydoped by the dopant compound generated in reaction zone 44 and thatmigrates into detection zone.

However, whereas the test device 10 of FIG. 1 relied upon measuring theinitial rate of interaction between the predetermined analyte and theoxidase enzyme included in reaction zone 14, the conductive sensor 40illustrated in FIG. 3 controls the amount of the dopant compound thatmigrates into the detection zone from reaction zone 44 by limiting themigration of the test sample 48 into the reaction zone 44 by asemipermeable membrane 42. By limiting the amount of the test sample 48,and therefore glucose, migrating through the restrictive semipermeablemembrane 42 to reaction zone 44, and by interacting and convertingsubstantially all of the glucose migrating into reaction zone 44 togenerate a dopant compound, the amount of glucose in the test sample 48is determined by correlating the rate of change in conductivity of thelayer or film of conducting polymer 46 in the detection zone to thediffusion rates of the test sample 48 through the semipermeable membrane42.

In particular, the test sample 48 including the predetermined analytecontacts the semipermeable membrane 42, and the test sample 48 permeatesthrough the semipermeable membrane 42 at a relatively slow, uniformrate. The permeation rate is a function of the analyte concentration ofinterest in the test sample 48. In effect, the migration of the testsample 48 through the semipermeable membrane 42 is the rate limitingstep of the interaction. When the test sample 48 contacts the reactionzone 44, an interaction between the predetermined analyte, like glucose,and the oxidase enzyme, like glucose oxidase, generates, either directlyor indirectly, a dopant compound, like molecular iodine. In thisparticular embodiment, and in contrast to the embodiment illustrated inFIG. 1, a sufficient amount of the oxidase and peroxidase enzymes areincorporated into the reaction zone 44 to insure that the permeationrate of glucose is the dominant limitation.

The dopant compound, like molecular iodine generated in the reactionzone 44, then migrates to the detection layer to oxidatively dope thelayer or film of conducting polymer 46 present in the detection zone ofthe conductive sensor 40. Accordingly, the layer or film of conductingpolymer 46 exhibits an increase in conductivity, with the increase inconductivity being related to the migration rate of the predeterminedanalyte through semipermeable membrane 42. The increase in conductivityis detected by a microelectrode assembly 50 that is in contact with thelayer or film of conducting polymer 46. Therefore, the concentration ofthe predetermined analyte is determined indirectly from the migrationrate of the predetermined analyte in the test sample 48 through thesemipermeable membrane 42.

An important feature of the preferred embodiment depicted in FIG. 3 isthe semipermeable membrane 42. The semipermeable membrane 42 hasphysical and chemical properties that allow the relatively slow, butuniform, migration of the predetermined analyte to the reaction zone 44.Furthermore, the permeation rate of ambient oxygen into, and through,the semipermeable membrane 42 is sufficiently high such that an oxygenlimitation on the enzymatic interactions occurring within the reagentzone 44 is precluded. In particular, the semipermeable membrane 42generally exhibits a diffusion constant for molecular oxygen of fromabout 5× 10⁻⁷ cm² /sec to about 5×10⁻⁶ cm² /sec, and for species such asglucose ranging from about 1×10⁻⁹ cm² /sec to about 5×8⁻⁸ cm² /sec.Preferably, the semipermeable membrane 42 is compatible with biologicalfluids, like blood and urine, and, to achieve the full advantage of thepresent invention, selectively screens assay interferents, likeascorbate and uric acid, from the test sample 48. Silicone-containingelastomers, applied as a film with a thickness of from about 3μ to about15μ, and preferably from about 5μ to about 10μ, are useful as thesemipermeable membrane 42. Other suitable semipermeable membranesinclude porous polypropylene, porous nylon, porous polycarbonate, porouspolyurethane, and similar porous materials; and combinations thereof. Toachieve the full advantage of the present invention, the semipermeablemembrane 42 is a film of from about 6μ to about 8μ in thickness.

In the preferred embodiment illustrated in FIG. 3, the semipermeablemembrane 42 retards the available glucose in the test sample 48 fromrapid contact with the reaction zone 44. It also should be understoodthat although ambient oxygen can permeate through the semipermeablemembrane 42 that the permeation is not sufficient to provide sufficientoxygen to interact with all the available glucose in the test sample 48.Therefore, the semipermeable membrane 42 retards the migration of theavailable glucose to the reaction zone 44 such that only about 3% orless of the available glucose reaches the reaction zone 44. Accordingly,the reaction zone 44 includes a sufficient amount of oxidase enzyme,peroxidase enzyme and ambient oxygen to interact with the glucose thatdoes migrate to the reaction zone 44. This interaction is detectedquickly, such as within 30 seconds, and preferably within 10 seconds,before continuing amounts of glucose migrate to the reaction zone 44 andencounter an oxygen limitation problem due to insufficient permeation ofambient oxygen to the reaction zone 44.

Therefore, the semipermeable membrane 42 has a sufficient thickness,like from about 3μ to about 15μ, to retard the migration of glucose tothe reaction zone 44. As the thickness of the semipermeable membrane 42increases, like above about 15μ, the time required to perform the assayis increased. Therefore, a thin semipermeable membrane 42 is preferred.

Therefore, the semipermeable membrane 42 preferably is a relatively thinfilm, of from about 6μ to about 8μ, that is permeable to both ambientoxygen and the predetermined analyte. A semipermeable membrane 42demonstrating a diffusion constant for oxygen of from about 5×10⁻⁷ cm²/sec to about 5×10⁻⁶ cm² /sec allows sufficient oxygen to enter thereaction zone 44 such that oxygen limitations are avoided. Furthermore,the semipermeable membrane 42 also should demonstrate a diffusionconstant for the predetermined analyte, like glucose, of from about1×10⁻⁹ cm² /sec to about 5×10⁻⁸ cm² /sec. A semipermeable membrane 42exhibiting this diffusion rate sufficiently retards the major amount ofthe predetermined analyte in test sample 48 from contacting the reactionzone 44, but permits a sufficient and known minor amount of thepredetermined analyte to contact the reaction zone 44 for a rapid andaccurate analyte determination. It is preferred that the semipermeablemembrane 42 be as thin as possible to allow migration of thepredetermined analyte and to eliminate interfering compounds.Accordingly, to provide a thin semipermeable membrane, the diffusionconstant the predetermined analyte through the semipermeable membraneshould be sufficient to allow an assay to be performed within 30seconds, and preferably within 10 seconds.

The semipermeable membrane 42 also provides the advantage of eliminatingthe effects of interfering compounds present in the test sample. Asnoted above, the generated molecular iodine can interact with commontest sample components, such as ascorbate, uric acid, acetaminophen, andsalicylates, before the molecular iodine can dope the conductingpolymer. This problem is partially resolved by utilizing a thin reactionzone 44 to generate the molecular iodine near the conducting polymersuch that doping of the conducting polymer occurs before the oxidationof interfering compounds. The semipermeable membrane 42 further resolvesthis problem by selectively screening anionic interfering compounds fromthe test sample 48. The anionic interfering compounds demonstrate adiffusion constant of only about 1×10⁻¹¹ cm² /sec to about 1×10⁻¹⁰ cm²/sec through the semipermeable membrane 42. Therefore, because thediffusion rate of the interfering compound through the semipermeablemembrane 42 is much slower than the diffusion rate of the predeterminedanalyte, the assay has been completed before an appreciable amount ofthe interfering compounds migrate to the reaction zone 44. Furthermore,the semipermeable membrane 42 effectively screens cellular material fromthe test sample 48, like the red blood cells in a whole blood sample,without clogging and further retarding the diffusion of thepredetermined analyte through the semipermeable membrane 42. Therefore,low molecular weight components in the blood plasma or blood serum canmigrate to the interaction layer 44.

In comparison to test device 10 of FIG. 1, the conductive sensor 40 ofFIG. 3 provides as fast an assay and substantially reduces the effectsof interfering compounds often found in the test sample. For example, ifthe semipermeable membrane 42 is a silicone-containing elastomer ofapproximately 6μ in thickness, the assay for the predetermined analytein the test sample 48 still can be completed within about 30 seconds,and usually within about 10 seconds. As will be demonstrated more fullyhereinafter, the semipermeable membrane 42 can be applied, reproducibly,as a thin, uniform layer, such that from the thickness of thesemipermeable membrane and the diffusion constant of the predeterminedanalyte through the semipermeable membrane, the amount of predeterminedanalyte that contacts the reaction zone 44 for interaction with theenzymes and generation of molecular iodine can be detected and measured.This measurement then can be correlated to the total amount of availablepredetermined analyte in the test sample.

To demonstrate the accuracy and sensitivity of a conductive sensor andthe method of the present invention, the following test device wasprepared and used in an assay for glucose. The test device included aconductive sensor 40 illustrated in FIG. 2. First, the microelectrodeassembly 50 of FIG. 3, and more fully illustrated as the microelectrodeassembly 30 in FIG. 2, was prepared. Accordingly, a wafer or layer of asmooth, nonconductive material, for example, but not limited to,silicon, ceramic, teflon, polycarbonate, polypropylene, kevlar,chrome-treated glass and glass, was used as the base 32 of themicroelectrode assembly 30 in FIG. 2 for the subsequent deposition ofthe interdigited patterns of conductive material 34 and 36.

Interdigited patterns of conductive material 34 and 36 are applied tothe top surface of base 32, and conductive contacting pads (not shown)for electrical connection to detection instrumentation are applied tothe bottom surface of base 32. The conductive contacting pads and theinterdigited patterns of conductive material 34 and 36 are usually ametal, and preferably are gold. Other suitable materials include silver,cermet, nickel and platinum In general, the only limitation on thematerial used as the conductive pads and the interdigited patterns ofconductive material 34 and 36 is that an ohmic contact exists betweenthe material and the layer or film of conducting polymer. Accordingly,aluminum is not a suitable material. The top and bottom surfaceconductors, i.e. the interdigited patterns of conductive material 34 and36 and the conductive contacting pads respectively, are connected byvias through the base 32. The interdigited patterns of conductivematerial 34 and 36 on the top surface of base 32 have a finger-likeappearance and have about a 10μ and about a 10μ gap and finger widths,respectively. In general, the gap widths range from about 10μ to about300μ, and the finger widths range from about 10μ to about 300μ. Fingerlengths range from about 100μ to about 400μ and usually are from about150μ to about 350μ. The total channel length ranges from about 100μ toabout 400μ and usually is from about 150μ to about 350μ.

Therefore, a test device of the present invention utilizes a smooth andnonconductive base for the microelectrode assembly, wherein the top, orsensing, surface of the microelectrode assembly includes an interdigitedmetal pattern, like a gold pattern, printed onto the base of themicroelectrode assembly by procedures well known in the art. Electricalcontact to the top sensing surface of the microelectrode assembly isaccomplished by gold vias to the back surface of the base of themicroelectrode assembly. Large gold contact pads positioned on the backsurface of the base of the microelectrode assembly provide an electricalconnection to detection instruments, such as a conductivity meter. Thisparticular configuration isolates the top, sensing surface of themicroelectrode assembly from the contact pads on the bottom surface ofthe base of the microelectrode assembly, and therefore avoids making anelectrical contact to detection instruments through the chemical layersof the reaction zone and detection zone that subsequently are positionedover the top sensing surface of the microelectrode assembly.

More particularly, on the top sensing surface of the electrode assemblytwo isolated halves of deposited metal on the surface form aninterdigited pattern. The overall dimensions of this interdigitedpattern ranged from 0.9 cm in diameter for gold printed on circularceramic to 687μ by 850μ for gold patterned on rectangular silicon. Thenumber of fingers protruding from each half was four and twentyrespectively. This undulating channel between each isolated half isgenerally 250μ wide and 33 mm long for the ceramic devices and 10μ wideand 27 mm long for the silicon devices. Finger widths are 250μ and 10μfor the ceramic and silicon devices, respectively. The fingers generallyextend about 300μ and are about 10μ wide. In accordance with animportant feature of the present invention, the test device measures theconduction of electrons from one half of the isolated interdigitedpattern to the other half of the pattern. Therefore, the size of the gapbetween each isolated half of the interdigited pattern is an importantfeature of the test device because the analyte assay is more sensitiveand more accurate as the gap is reduced.

Accordingly, after manufacturing a microelectrode assembly, a layer orfilm of conducting polymer, from about 100 Å to about 10,000 Å inthickness, is deposited in the gap between the isolated interdigitedpatterns of conductive material. Then a reaction zone, from about 0.1μto about 10μ in thickness, is deposited over the layer of conductingpolymer, such that a dopant compound is produced upon contact betweenthe reaction zone and a test sample, and dopes the conducting polymer.Above the reaction zone is deposited a semipermeable membrane of about3μ to about 15μ in thickness.

Therefore, onto the top sensing surface of the microelectrode assemblyis positioned a layer or film of a suitable conducting polymer. Asstated previously, a suitable conducting polymer provides a resistanceupon exposure to molecular iodine in the range from about 10¹ ohm toabout 10⁹ ohm when the polymer film has a thickness of from about 100 Åto 1000 Å. Furthermore, the conducting polymer should be stable and easyto process such that the conducting polymer can be applied to themicroelectrode assembly as a uniform, thin film or layer by spincoating, film casting, jet printing or a similar application techniqueknown in the art. The thickness of the film or layer of the conductingpolymer is sufficient to fill the gaps between the two isolatedinterdigited patterns on the substrate electronic template and tocompletely coat, or cover, the interdigited patterns such that theinterdigited patterns of conducting material do not contact the reactionzone of the conductive sensor.

Accordingly, suitable conducting polymers include, but are not limitedto, the polythiophenes, polypyrroles, polyfurans, poly(thienylenevinylenes), and poly(furylene vinylenes) that demonstrate sufficientsolubility in organic solvents to be processed into a film. Preferably,the conducting polymer is a polythiophene, such as apoly(3-alkylthiophene), wherein the alkyl group includes from about fourto about twenty, and preferably from about six to about twelve, carbonatoms. To achieve the full advantage of the present invention, the alkylgroup of the poly(3-alkylthiophene) includes from about six to aboutnine carbon atoms. These particular polythiophenes have demonstrated asufficient electrical conductivity and satisfactory processingproperties, such as solubility in organic solvents, to provide a thin,uniform layer or film of conducting polymer on the microelectrodeassembly. As previously discussed, the conducting polymer can be admixedwith nonconducting polymers, or the conducting polymer can be acopolymer, to improve the processing, physical and morphologicalproperties of the film or layer of the conducting polymer.

Therefore, in the manufacture of a conductive sensor 40 of FIG. 3, aparticularly useful conducting polymer, poly(3-octylthiophene), firstwas dissolved in a suitable solvent, like xylene, at a concentration of5 mg (milligrams) of conducting polymer per ml (milliliter) of solvent.In general, it has been found that the concentration of the conductingpolymer in the solvent should range from about 2 mg/ml to about 15mg/ml, and preferably from about 3 mg/ml to about 10 mg/ml. Moreconcentrated solutions of the conducting polymer also can be used aslong as the viscosity of the solution is suitable for casting a uniform,thin film of conducting polymer.

Other suitable solvents that can be used in place of, or in addition to,xylene include, but are not limited to, benzene, toluene, chloroform,methylene chloride, trichloroethylene, tetrahydrofuran,1,2-dichloroethane, nitrobenzene, dimethylsulfoxide, dimethylformamide;and combinations thereof. In general, any organic solvent capable ofsolubilizing the conducting polymer and having a sufficient vaporpressure to evaporate from the film of conducting polymer is suitablefor use in the present invention.

Before applying the solution of the conducting polymer to themicroelectrode assembly, the microelectrode assembly is cleaned with asolvent, like chloroform, by washing the microelectrode assembly, thenspinning the electronic template dry on a PHOTO-RESIST SPINNER,available from Headway Research Inc., Garland, Tex. With themicroelectrode assembly at rest on the spinner, the top sensing surfaceof the microelectrode assembly is flooded with the conductingpolymer-xylene solution and then spun to dryness at 3000 rpm forapproximately 20 seconds on the PHOTO-RESIST SPINNER to produce a thin,uniform film. The conducting polymer film then is allowed to air dry forat least 30 minutes. The dried layer, or film, of conducting polymer hasa thickness of approximately 200 Å to approximately 300 Å.

After the layer or film of conducting polymer has dried sufficiently,the reaction zone is positioned in laminar contact with the top surfaceof the layer or film conducting polymer. The reaction zone comprises thehydratable host matrix homogeneously incorporating a suitable oxidaseenzyme and other enzymes and dopant compound precursors to interact withthe predetermined analyte and to generate a dopant compound. Thereaction zone is sufficiently thick such that a sufficient amount ofenzymes and dopant compound precursors can be included in the reactionzone, and is sufficiently thin such that the dopant compound isgenerated near the layer or film of conducting polymer to dope theconducting polymer before the dopant compound undergoes an interactionwith an interfering compound present in the test sample.

Accordingly, in the preparation and application of the reaction zone tothe test device, the following stock gelatin solution (20% W/W) wasprepared by adding about 140 g (grams) of gelatin, such as GELATINZKN-407, available from Sigma Chemical Co., St. Louis, Mo., in 560 g ofdistilled water. After allowing the gelatin to swell in the waterapproximately 15 to approximately 20 minutes, the gelatin-water mixturewas heated to about 45° C. to melt the gelatin. After approximately oneto 3 hours, the gelatin was melted and the pH of the gelatin-watermixture was adjusted to 7.0 with sodium hydroxide. The stock gelatinsolution then was stored at a temperature of from 0° C. to 5° C. In thisparticular embodiment, gelatin was used as the hydratable host matrix ofthe reaction zone. However, other suitable materials useful as thehydratable host matrix include, but are not limited to, collagen, silkfibroin, crosslinked albumin, a polyacrylamide andpoly(2-hydroxyethylmethacrylate) and combinations thereof.

After forming the stock gelatin solution, a gelatin casting compositionwas prepared. This particular casting composition included:

    ______________________________________                                        Gelatin (20% W/W)   50.00%  (by weight)                                       BES buffer 1M, pH 7                                                                               5.00%                                                     GOD                 0.11%                                                     POD                 0.44%                                                     Distilled Water     44.45%                                                                       100.00%,                                                   ______________________________________                                    

wherein the gelatin is the stock gelatin solution described above; BESis a buffer solution including N,N-bis(2-hydroxyethyl)-2-aminoethanesulfonic acid and sodium N,N-bis(2-hydroxyethyl)-2-aminoethanesulfonate; GOD is glucose oxidase having an activity of at least about85 U/mg (units per milligram, wherein a unit is defined as the abilityto oxidize 1.0 μmole (micromole) of β-D-glucose to D-gluconic acid andhydrogen peroxide per minute at pH 5.1 at 35° C.); and POD is aperoxidase enzyme having an activity of approximately 1160 U/mg (unitsper milligram, wherein a unit is defined as the ability to form one mgpurpurogallin in 20 seconds from pyrogallol at pH 6.0 and 20° C.) and anRZ of at least 3, wherein the term RZ is a measure of hemin content andcan be indicative of enzyme purity. Suitable peroxidase enzymes include,but are not limited to, horseradish peroxidase, lactoperoxidase,microperoxidase or combinations thereof. However, it should beunderstood that other compounds that have peroxidase activity, i.e.behave like peroxidase in the presence of hydrogen peroxide, also can beincluded in the hydratable host matrix in addition to or in place of theperoxidase. Compounds exhibiting peroxidase activity that can beincluded in the hydratable host matrix include, but are not limited to,molybdenum(VI) transition metal catalysts and similar transition metalcatalysts. Molybdenum(VI) transition metal catalysts are known tointeract with hydrogen peroxide to convert iodide ion to moleculariodine. Furthermore, other suitable buffers well-known in the art, suchas phosphate buffers, also can be used in the casting solution.

It should be understood that the above gelatin casting composition wasused when glucose was the predetermined analyte of interest However, ifa different analyte is of interest, the GOD is replaced by theappropriate oxidase enzyme that interacts with that particularpredetermined analyte. For example, if alcohol is the predeterminedanalyte of interest, a sufficient amount of alcohol oxidase replaces theGOD.

The gelatin casting composition was prepared by melting 50 g of the 20%stock gelatin solution in a 45° C. water bath. In a separate beaker, thebuffer and distilled water were premixed to form an aqueous buffersolution. The GOD and POD then were added to the aqueous buffersolution. After the enzymes were dissolved in the aqueous buffersolution, and after the stock gelatin solution was melted, the twosolutions were combined and the resulting mixture was stirred gently at40°-45° C. for at least 15 minutes.

The above-described gelatin casting composition illustrates a castingcomposition that provides a reagent zone of a conductive sensor of thepresent invention. In general, it has been found that a suitable castingcomposition comprises:

    ______________________________________                                        Hydratable host matrix                                                                          2% to 20% (dry basis) (by                                                     weight)                                                     Buffer            q.s. to sufficiently                                                          buffer to about pH 7.sup.1)                                 Oxidase Enzyme    0.05% to 1.5%                                               Peroxidase Enzyme 0.4% to 0.6%                                                Dopant Compound   0% to 2%                                                    Precursor                                                                     Water             q.s. to 100%                                                ______________________________________                                         .sup.1) For example, 1M BES buffer, 4% to 6% by weight provide pH 7.     

1) For example, 1 M BES buffer, 4% to 6% by weight to provide pH 7.

In addition to preparing the casting composition, a crosslinkingcomposition including the following ingredients was prepared by simplyadmixing the composition ingredients, and stirring, to provide ahomogeneous crosslinking composition. The composition pH then wasadjusted to about 7.

    ______________________________________                                        EDAC                   2.0% (by weight)                                       SURFACTANT 10 G (5% aqueous                                                                          8.0%                                                   solution)                                                                     BES buffer 1M, pH 7    2.5%                                                   Distilled Water        87.5%                                                                        100.0%                                                  ______________________________________                                    

In the above crosslinking composition, EDAC is1-ethyl-3-(3-dimethylaminopropyl)carbodiimide, available from SigmaChemical Co., St. Louis, Mo., and used to harden and crosslink thegelatin-based hydratable host matrix of the reaction zone. Othersuitable crosslinking agents include, but are not limited to, thecrosslinking reagents listed in the 1987 catalog of Sigma Chemical Co.,St. Louis, Mo., on page 452. SURFACTANT 10G is anonylphenol-polyglycidol surfactant, available from Olin Chemical,Stamford, Conn. This optional surfactant ingredient is included toimprove the overall coating properties of the gelatin-based film. Ingeneral, a suitable crosslinking composition comprises from about 1% toabout 5%, and preferably from about 2% to about 3%, by weight of acrosslinking agent; from 0% to about 1%, and preferably from about 0.3%to about 0.4%, by weight of a surfactant; a sufficient amount of abuffer to maintain a pH of about 7; and water.

The reaction zone was incorporated onto the test device by firstapplying the casting composition and then the crosslinking compositionto the test device with a PHOTO-RESIST SPINNER. The PHOTO-RESISTSPINNER, or an equivalent device, reproducibly applies a thin, uniformreaction zone film. First, an aliquot of the gelatin casting solutionwas applied over the layer or film of conducting polymer. The spinnerwas rapidly accelerated to approximately 2000 rpm, and maintained atapproximately 2000 rpm for approximately 20 seconds. The resulting wetfilm of gelatin casting solution was allowed to air dry forapproximately 15 minutes, and the film of gelatin casting compositionhad hardened sufficiently to allow application of the crosslinkingcomposition. The crosslinking composition then was applied over theentire wet gelatin film surface. After a short contact time between thecrosslinking composition and the wet film of gelatin castingcomposition, the test device again was spun on the PHOTO-RESIST SPINNER.Then, the test device was allowed to air dry for at least about 1 hourbefore use. The dried, crosslinked gelatin-based reaction zone film hada thickness of from approximately 1μ to approximately 3μ.

Then, positioned in laminar contact with the top surface of the reactionzone, a thin film of a silicone elastomer was applied to the conductivesensor. The silicone elastomer served as the semipermeable membrane. Thesemipermeable membrane was applied to the conductive sensor by the samebatch type processes described above to provide, reproducibly, a thin,uniform film of silicone elastomer having a thickness of about 6μ toabout 8μ. It should be understood that for each of the three layers orfilms, i.e. the conducting polymer, the reaction zone and thesemipermeable membrane, once the desired thickness of the film isdetermined, the semiconductor processing techniques utilized tomanufacture the conductive sensors provide layers or films that arereproducible to within ±5% in thickness. It is important that thethickness of each layer be controlled such that the thickness of theindividual layers is essentially eliminated as a variable parameter inthe assay for a predetermined analyte. Therefore, in accordance with animportant feature of the present invention, the thickness of each layeris controlled to within ±5% of the desired, predetermined thickness.

The ability to reproducibly provide thin films of uniform thickness isan important feature of the present invention. For example, the abilityto provide uniform, thin films allows the economical and facilemanufacture of miniaturized test devices. Therefore, extremely smalltest samples can be used, such as from about 0.1 μL to about 5 μL. Incontrast, present test devices require at least about 5 μL, and up toabout 20 μL of test sample to provide an accurate assay. Theselarger-sized blood samples require invasive and more painful test sampleacquisition techniques. The miniature size of the test device alsoallows the economical manufacture of a disposable test device.

In addition to allowing the easy and economical manufacture ofdisposable test devices that require very small sample sizes, theability to reproducibly cast uniform thin films provides a fast andaccurate assay for a predetermined analyte. The thin films allow asufficient amount of the predetermined analyte to migrate through thesemipermeable membrane to contact the reaction zone and interact withthe enzymes and the dopant compound precursor. After interacting togenerate the dopant compound, the conducting polymer then is doped toprovide an assay within about 30 seconds, and preferably within about 10seconds. In addition, the assay is accurate and sensitive because thefilms or layers are sufficiently thin such that the dopant compound isgenerated in proximity to the conducting polymer layer. Accordingly, thedopant compound can dope the conducting polymer before undergoingreactions with interferent compounds often present in the test sample.Therefore, essentially all the dopant compound is available to dope thepolymer to provide an accurate assay. Furthermore, because theconducting polymer is thin, the sensitivity of the conductive sensor isincreased because a relatively large change in conductivity is observedfor a relatively small amount of generated dopant compound.

The ability to reproducibly provide thin uniform films also is importantbecause only a small portion of the total available predeterminedanalyte in the test sample is interacted to generate the dopantcompound. Therefore, to accurately correlate the change in conductivityof the conducting polymer to the total amount of predetermined analytein the test sample, the layers in the conductive sensor are manufacturedat a reproducible thickness such that assay measurements are indicativeof the actual concentration of predetermined analyte in the test sample,and not of the apparent concentration of predetermined analyte in thetest sample because of appreciable thickness variances in any of thethree films or layers in the sensor.

In accordance with an important feature of the present invention, a testdevice prepared by the above described method, but lacking thesemipermeable membrane, i.e. a test device 10 of FIG. 1, accuratelyassayed standardized glucose solutions including from 25 mg/dL to 500mg/dL glucose. The standardized glucose solutions further includediodide ions. As discussed above, in a test device of the presentinvention however, the iodide ions are included in the reaction zone ofthe conductive sensor in an amount ranging from about 30 mM to about 500mM. The conductivity of the film or layer of conducting polymer in thedetection zone was measured at an applied voltage of 0.1 V (volt), andelectrical currents the range of from 0.1 μamp (microamp) to 5 μamp werefound as typical.

As will be demonstrated more fully hereinafter in the detailedexplanation of FIGS. 4 through 7, fully-assembled test devices weretested by applying a 0.1 volt potential across the two isolatedinterdigited electrodes in the microelectrode assembly and measuring theincrease in conductivity of the film or layer of conducting polymer asthe conducting polymer is doped by the dopant compound generated in thereaction zone. It was found that a linear change in conductivity of thefilm or layer of conducting polymer results from the enzymatic oxidationof glucose in standardized glucose solutions including from 25 to 500mg/dL glucose and of 30 to 50 mM/L potassium iodide. Ultimately, theiodide ion was oxidized to molecular iodine that dopes the film or layerof conducting polymer, and accordingly, increases the conductivity ofthe conducting polymer film such that currents in the range of from 1 to5 μamps were observed. In general, the plots presented in FIG. 4illustrate the results for glucose assays of standardized glucosesolutions. The plots in FIGS. 6 and the plot in FIG. 7 were identicallyderived, except the glucose oxidase activity in the reaction zone wasincreased slightly from 85 U/mg to 118 U/mg of casting solution. InFIGS. 5 and 7, a dose response plot is generated in each experiment bypolling the current at 10 seconds.

In particular, the dose response plots graphed in FIG. 4 illustrate theresults of assays for glucose utilizing the method and device of thepresent invention. Standardized glucose solutions including from 25mg/dL to 500 mg/dL, and further including 30 mM potassium iodide, wereassayed by a test device 10 as illustrated in FIG. 1. Initially, it canbe observed that the greater the concentration of glucose in solution,the greater the initial change in conductivity, and the greater thetotal change in conductivity, of the film or layer of conductingpolymer. In addition, it also is observed that the change inconductivity of the film or layer of conducting polymer is greatest whenthe percentage amount of dopant compound in the conducting polymer layeris relatively low, i.e. at the early stages of the glucose-glucoseoxidase interaction, such as within about 30 seconds, and especiallywithin about 15 seconds, of the onset the glucose-glucose oxidaseinteraction. As a result, the supply of ambient oxygen has not beensignificantly depleted.

Therefore, the most sensitive measurement for the change in conductivityof the conducting polymer film is made within approximately the first 30seconds, and preferably within approximately the first 10 seconds, afterthe test sample contacts the reaction zone of the test device. Thechange in conductivity after the first 25 seconds does not appreciablyincrease because of ambient oxygen supply depletion or because theamount of glucose in the reaction zone, usually less than about 1% ofthe glucose in the test sample, has interacted with the enzymes and thedopant compound precursor.

It should be understood that only a small amount of the availableglucose present in the test sample is actually converted to eventuallygenerate molecular iodine, and that the amount of glucose in the testsample is determined from the initial reaction rate between the glucoseand the enzymes to generate molecular iodine. In addition, in acommercial test device, the iodide ion is included in the reaction zonewith the enzymes and other necessary reagents rather than adding theiodide ion to the glucose-containing test sample.

The graph plotted in FIG. 5 illustrates the linear response of thechange in conductivity of the layer of conducting polymer to the amountof glucose in the test sample. In each assay, the conductivity of thelayer of conducting polymer was measured 10 seconds after theglucose-containing sample contacted the reaction zone of the testdevice. The linear relationship illustrated in FIG. 5 shows that a testsample including an unknown amount of glucose can be assayed by thedevice and method of the present invention. For example, the small testsample is introduced into the test device. Then, ten seconds after thetest sample contacts the reaction zone, the conductivity of the layer ofconducting polymer is measured. From a linear graph as presented in FIG.5 and derived from standardized glucose solutions, the measuredconductivity of the layer of conducting polymer can be correlated to theamount of glucose in the test sample. It has been found that the mostsensitive and accurate assays are achieved when the conductivity ismeasured from about 5 seconds to about 30 seconds after the test samplecontacts the reaction zone.

The plots in FIGS. 6 and 7 show essentially identical results obtainedwhen the amount of glucose oxidase in the casting solution, and hence inthe reaction zone is increased. The greatest change in conductivity isobserved in the first 30 seconds, and especially in the first 15seconds, of the reaction zone interaction (FIG. 6) and the linearity ofthe measured conductivity of the layer of conducting polymer to theconcentration of glucose in the test sample is maintained (FIG. 7).Hence the method and device of the present invention provide a fast,economical, sensitive and accurate assay of a liquid test sampleincluding glucose in amounts as low as 0 mg/dL and up to 600 mg/dL.

In accordance with another important feature of the present invention,another test device was prepared by the above-described method andincluded a semipermeable membrane, i.e. a test device 40 of FIG. 3. Thistest device accurately assayed standardized glucose solutions includingfrom 5 mg/dL to 500 mg/dL glucose. The standardized glucose solutionsfurther included iodide ions. As discussed above and illustrated below,in a test device of the present invention, the iodide ions preferablyare included in the reaction zone of the conductive sensor in an amountranging from about 50 mM to about 500 mM. The conductivity of the filmor layer of conducting polymer in the detection zone was measured at anapplied voltage of 0.1 V (volt), and electrical currents the range offrom 0.1 μamp (microamp) to 5 μamp were found as typical.

As will be demonstrated more fully hereinafter in the detailedexplanation of FIGS. 8 and 9, fully-assembled test devices were testedby applying a 0.1 volt potential across the two isolated interdigitedelectrodes in the microelectrode assembly and measuring the increase inconductivity of the film or layer of conducting polymer as theconducting polymer is doped by the dopant compound generated in thereaction zone. It was found that a linear change in conductivity of thefilm or layer of conducting polymer results from the enzymatic oxidationof glucose in standardized glucose solutions including from 5 mg/dL to500 mg/dL glucose and 150 mM/L potassium iodide. Ultimately, the iodideion was oxidized to molecular iodine that dopes the film or layer ofconducting polymer, and accordingly, increases the conductivity of theconducting polymer film such that currents in the range of from 1 to 5μamps were observed.

In particular, in these assays, the test device included a conductivesensor 40 wherein the reaction zone 44 included chitosan as thehydratable matrix material. In this embodiment, a 3% by weight aqueoussolution of shellfish chitosan, available from PROTAN, INC., Commack,N.Y., first was prepared. Then a sufficient amount of glucose oxidaseand peroxidase was added to the 3% chitosan solution to provide achitosan solution that includes 10 mg/ml (milligrams per milliliter)each of glucose oxidase and perioxidase. This chitosan solution then wascast over a layer of polyoctylthiophene, as previously described.

After allowing chitosan layer to dry sufficiently to provide thereaction zone 44, a semipermeable membrane 42 was cast over the chitosanlayer. The semipermeable membrane 42 was applied by casting an aqueoussolution including about 75% by weight of the elastomer, Dow CorningLatex 3-5035, available from Dow Corning Corporation, Midland, Mich.,and about 0.1% by weight OLIN SURFACTANT 10G onto the chitosan-basedreaction zone 44. After drying the film of elastomer-based semipermeablemembrane 42, a conductive sensor 40 was provided to assay a test samplefor glucose.

In general, the plots presented in FIGS. 8 and 9 illustrate the resultsof glucose assays of standardized glucose solutions using a conductivesensor 40 of FIG. 3. In both FIGS., a dose response plot is generated ineach experiment by polling the current at 10 seconds.

In particular, the dose response plots graphed in FIG. 8 illustrate theresults of assays for glucose on standardized glucose solutionsincluding from 5 mg/dL to 500 mg/dL glucose, and further including 30 mMpotassium iodide. Initially, it can be observed that the greater theconcentration of glucose in solution, the greater the initial change inconductivity, and the greater the total change in conductivity, of thefilm or layer of conducting polymer. In addition, it also is observedthat the change in conductivity of the film or layer of conductingpolymer is greatest when the percentage amount of dopant compound in theconducting polymer layer is relatively low, i.e. at the early stages ofthe glucose-glucose oxidase interaction, such as within about 25seconds, and especially within about 12 seconds, of the onset theglucose-glucose oxidase interaction. Therefore, the most sensitivemeasurement for the change in conductivity of the conducting polymerfilm is made within approximately the first 30 seconds, and preferablywithin approximately the first 10 seconds, after the test samplecontacts the reaction zone of the test device.

The graph plotted in FIG. 9 illustrates the linear response of thechange in conductivity of the layer of conducting polymer to the amountof glucose in the test sample. In each assay, the conductivity of thelayer of conducting polymer was measured 10 seconds after theglucose-containing sample contacted the reaction zone of the testdevice. The linear relationship illustrated in FIG. 9 shows that a testsample including an unknown amount of glucose can be assayed by thedevice and method of the present invention.

Hence the device of the present invention depicted in FIG. 3 provides afast, economical, sensitive and accurate assay of a liquid test sampleincluding glucose in amounts as low as 0 mg/dL and up to 600 mg/dL.

To demonstrate that a conductive sensor of the present invention canaccurately assay a test sample for glucose when the iodide ion isincluded in the reaction zone of the conductive sensor the followingassays were performed using a test device including a conductive sensor40 as depicted in FIG. 3. It has been found that the iodide ion is mosteasily incorporated into the reaction zone 44 in the form of atetraalkylammonium iodide, wherein the alkyl group includes from one toabout four carbon atoms. Preferably, the tetraalkylammonium iodide istetraethylammonium iodide. However, the tetramethyl, tetrapropyl andtetrabutyl iodides also can be used. Similarly, potassium iodide,lithium iodide, sodium iodide and other water-soluble iodide salts canbe used as the source of iodide ions in the reaction zone. The iodidesalt is included in the casting solution used to form the reaction zonein a concentration ranging from about 50 mM to about 500 mM, andpreferably in the range of from about 75 mM to about 300 mM.

FIG. 10 includes the dose response plots for glucose assays utilizing aconductive sensor 40 of FIG. 3 wherein tetraethylammonium iodide wasincluded in a chitosan-based reaction zone 44 at a concentration ofabout 100 mM. In this experiment, the conductive sensor was slightlymodified in that a second chitosan-based film, including thetetraethylammonium iodide, was applied over a first chitosan-based filmthat included the glucose oxidase and the peroxidase. Although thisparticular reaction 44 included two distinct chitosan films, the iodidesalt, glucose oxidase and peroxidase can be included in a singlechitosan, or gelatin, film. A semipermeable membrane 42, comprising asilicone-based elastomer, was cast over the chitosan-based filmincluding the tetraethylammonium iodide. The above-described test devicewas used to assay standardized solutions including 100 mg/dL, 200 mg/dLand 400 mg/dL of glucose. The results of the glucose assay areillustrated in FIG. 10, showing that the device analyzes for glucosewithin about 15 seconds.

Several enzyme-based assays involve the generation of hydrogen peroxideas a reaction product. As described above for glucose, the generation ofhydrogen peroxide in the glucose-glucose oxidase interaction provides asubstrate for the peroxidase enzyme to generate the molecular iodinedopant compound. The molecular iodine then dopes the film or layer ofconducting polymer, and the amount of glucose in the test sample isdetermined from the change in conductivity of the conducting polymerlayer. However, in addition to glucose oxidase, oxidase enzymes thatemploy oxygen as a mediator, and therefore produce hydrogen peroxideupon interaction with the appropriate substrate include, but are notlimited to:

cholesterol oxidase,

aryl-alcohol oxidase,

L-gluconolactone oxidase,

galactose oxidase,

pyranose oxidase,

L-sorbase oxidase,

pyridoxin 4-oxidase,

alcohol oxidase,

L-2-hydroxyacid oxidase,

pyruvate oxidase,

oxalate oxidase,

glyoxylate oxidase,

dihydro-orotate oxidase,

lathosterol oxidase,

choline oxidase,

glycolate oxidase,

glycerol-3-phosphate oxidase,

xanthine oxidase,

sarcosine oxidase,

N-methylamino-acid oxidase,

N⁶ -methyl-lysine oxidase,

6-hydroxyl-L-nicotine oxidase,

6-hydroxy-D-nicotine oxidase,

nitroethane oxidase,

sulphite oxidase,

thiol oxidase,

cytochrome c oxidase,

Pseudomonas cytochrome oxidase,

ascorbate oxidase,

o-aminophenol oxidase, and

3-hydroxyanthranilate oxidase.

As a result, in accordance with the device and method of the presentinvention, conductive sensors for a particular predetermined analyte canbe designed utilizing the appropriate oxidase enzyme in the same deviceand method described above the assay of glucose. However, it should beunderstood, that for the embodiment illustrated in FIG. 3, thesemipermeable membrane can inhibit the passage of certain highermolecular weight compounds, like cholesterol, through the semipermeablemembrane. Therefore, the conductive sensor of the present inventionembodied in FIG. 3 may not be suitable in assays for higher molecularweight analytes, like cholesterol. In general however, as long as thepredetermined analyte exhibits a diffusion constant of at least about1×10⁻⁹ cm² /sec through the semipermeable membrane, then the presence orconcentration of the predetermined analyte can be determined by themethod and device of the present invention.

The primary features relating to the test device and method of thepresent invention have been repeatedly observed. The new and unexpectedresults arising from the method of the present invention provides testdevices designed to assay a liquid test sample for a predeterminedanalyte capable of interacting with oxygen and an oxidase enzyme.

From the foregoing, it is seen that the present invention is welladapted to attain all of the objects hereinabove set forth. The methodand device have the advantages of convenience, simplicity, relativeeconomy, disposability, sensitivity and accuracy. Among the advantagesof the present invention is that the device operates nonoptically; hasexcellent shelf life stability; can be constructed at relatively lowcost; can be reproducibly manufactured by semiconductor processingtechniques; has a great degree of flexibility with respect to format;requires a small, noninvasive amount of sample; and can be constructedto have a relatively small size.

For example, it is envisioned that a test device of the presentinvention is an economical, miniaturized, disposable device. Eachcomponent in every embodiment of the present invention can bemanufactured in a batch processing technique. In addition, themicroelectrode assembly can be manufactured by a number of techniqueswell-known in the art utilizing a silicon, ceramic, glass or plasticbase. Furthermore, the conductive sensor is stable because the methodand device utilize an undoped, reduced layer conducting polymer, asopposed to a less stable, oxidized conducting polymer utilized in mostprior art conductive sensors.

Although the present invention is primarily directed to assaying liquidmedia for various clinically significant substances or constituents inbiological fluids, such as urine and blood, including lysed or unlysedblood, blood plasma and blood serum, it should be understood that thedevice and method of the present invention are useful for assays ofnonbiological fluids, including swimming pool water, wines, etc.

It will be understood that the present disclosure has been made only byway of preferred embodiment and that numerous changes in details ofconstruction, combination, and arrangement of parts can be resorted towithout departing from the spirit and scope of the invention ashereunder claimed.

We claim:
 1. A conductive sensor for assaying a test sample for thepresence or concentration of glucose, said conductive sensorcomprising:a) a semipermeable membrane capable of effectively separatingcellular material and interfering components from a test sample andcapable of allowing glucose to permeate through the semipermeablemembrane at a uniform rate said semipermeable membrane having adiffusion constant for glucose in the range of from about 1×10⁻⁹ cm²/sec to about 5×10⁻⁸ cm² /sec and a diffusion constant for molecularoxygen in the range of from about 5×10⁻⁷ cm² /sec to about 5×10⁻⁶ cm²/sec and having a thickness of from about 3μ to about 15μ; b) a layer ofa host matrix in contact with the semipermeable membrane and permeableto the glucose, said hose matrix layer having homogeneously incorporatedtherein glucose oxidase, a compound having peroxidase activity and adopant compound precursor, wherein the glucose, the glucose oxidase, thecompound having peroxidase activity and the dopant compound precursorinteract to form a dopant compound; c) a layer of a polymer in itsreduced form which demonstrates an increase in electrical conductivitywhen it is converted to its oxidized form in contact with the hostmatrix layer such that at least a portion of the dopant compoundgenerated in the host matrix layer migrates to and oxidatively dopes thepolymer layer to thereby increase the conductivity of this layer; and d)means operatively connected to the polymer layer for measuring a changein conductivity of polymer layer.
 2. The conductive sensor of claim 1wherein the means for measuring the change in conductivity of thepolymer layer comprises a microelectrode assembly in contact with thepolymer layer, said microelectrode assembly constructed to sense thechange in conductivity of the polymer layer in response to the oxidativedoping of the polymer layer by the dopant compound.
 3. The conductivesensor of claim 2 wherein the microelectrode assembly comprises aninterdigited pair of metal electrodes having an insulating spacing offrom about 10μ to about 300μ.
 4. The conductive sensor of claim 1wherein the compound having peroxidase activity is horseradishperoxidase, lactoperoxidase, microperoxidase or combinations thereof. 5.The conductive sensor of claim 1 wherein the host matrix layer has athickness of from about 0.1μ to about 10μ and comprises gelatin,chitosan, silk fibroin, collagen, poly(2-hydroxyethylmethacrylate),polyacrylamide or combinations thereof.
 6. The conductive sensor ofclaim 1 wherein the dopant compound precursor is iodide ion, said iodideion incorporated into the host matrix layer as an iodide salt selectedfrom the group consisting of lithium iodide, sodium iodide, potassiumiodide, a tetraalkylammonium iodide wherein the alkyl group includesfrom one to about four carbon atoms, and combinations thereof.
 7. Theconductive sensor of claim 6 wherein the compound having peroxidaseactivity is a molybdenum(VI) transition metal catalyst.
 8. Theconductive sensor of claim 6 wherein the polymer layer has a thicknessof less than 10,000 Å and comprises a poly(alkylthiophene), apolythiophene, a poly(thienylene vinylene), a poly(furylene vinylene), apolypyrrole, a polyfuran, a polyaniline or a combination thereof.
 9. Theconductive sensor of claim 8 wherein the poly(alkylthiophene) is apoly(3-alkylthiophene) wherein the alkyl group includes from about fourcarbon atoms to about 20 carbon atoms.
 10. The conductive sensor ofclaim 1 wherein the semipermeable membrane has a thickness in the rangeof from about 5μ to about 10μ.
 11. The conductive sensor of claim 1wherein the semipermeable membrane has a thickness in the range of fromabout 6μ to about 8μ.
 12. The conductive sensor of claim 1 wherein thesemipermeable membrane is an elastomeric compound.
 13. The conductivesensor of claim 1 wherein the semipermeable membrane is asilicone-containing elastomer.
 14. The conductive sensor of claim 1wherein the semipermeable membrane comprises polypropylene, nylon,polycarbonate, polyurethane or a combination thereof.
 15. The conductivesensor of claim 1 wherein the glucose oxidase and compound peroxidasehaving activity are in a first matrix layer and the dopant compoundprecursor is in a second matrix layer immediately adjacent to the firstmatrix layer.
 16. The conductive sensor of claim 15 wherein the firstmatrix layer and the second matrix layer are comprised of chitosan. 17.The conductive sensor of claim 16 wherein the dopant compound precursoris tetraethylammonium iodide.
 18. The conductive sensor of claim 1wherein the host matrix comprises a first layer containing the glucoseoxidase and compound peroxidase having activity and a second layerimmediately adjacent to the first layer which contains the dopantcompound precursor material.
 19. The conductive sensor of claim 18wherein the first and second layers of the host matrix are comprised ofchitosan.
 20. The conductive sensor of claim 19 wherein the dopantcompound precursor is tetraethylammonium iodide.
 21. A conductive sensorfor assaying a biological fluid for the presence or concentration of apredetermined analyte, said predetermined analyte capable of interactingwith an oxidase enzyme, and said conductive sensor comprising:a) anelastomer-based semipermeable membrane having a thickness of from about5μ to about 10μ, said elastomer-based semipermeable membrane capable ofeffectively separating cellular material and interfering components froma biological fluid and capable of allowing a predetermined analyte topermeate through the elastomer-based semipermeable membrane at a uniformrate; b) a layer of a host matrix, comprising gelatin, chitosan or acombination thereof and having a thickness of from about 0.2μ to about5μ, in contact with the semipermeable membrane, said host matrix layerpermeable to the predetermined analyte and said host matrix layer havinghomogeneously incorporated therein an oxidase enzyme capable ofinteracting with the predetermined analyte, a peroxidase enzyme or amolybdenum(VI) transition metal catalyst, and iodide ion, wherein thepredetermined analyte, the oxidase enzyme, the peroxidase enzyme or themolybdenum(VI) catalyst, and the iodide ion interact to generatemolecular iodine; c) a layer of a conducting polymer having a thicknessof from about 100 Å to about 2000 Å in contact with the host matrixlayer such that at least a portion of the molecular iodine generated inthe host matrix layer migrates to and oxidatively dopes the conductingpolymer layer; and d) a microelectrode assembly in contact with theconducting polymer layer, said microelectrode assembly adapted to sensea change in conductivity of the conducting polymer layer in response tothe oxidative doping of the conducting polymer layer by the moleculariodine.